Pressure sensors for small-scale applications and related methods

ABSTRACT

Certain pressure sensor devices are much smaller than prior art devices, yet are at least as sensitive as the prior art devices. A capacitive pressure sensor can include a flexible substrate that permits bending of a pressure sensing region without significantly affecting operation thereof. The pressure sensor can include a flexible membrane in which an electrode is sandwiched between two layers of polymeric material. The sandwiched electrode can be extremely close to a reference electrode so as to provide for highly sensitive capacitance readings, yet the membrane can be restricted from contacting the reference electrode under high pressure condition.

CROSS REFERENCE TO RELATED APPLICATIONS

This application claims the benefit of U.S. Provisional PatentApplication Nos. 61/546324 filed Oct. 12, 2011, 61/660402 filed Jun. 15,2012, and 61/712,579 filed Oct. 11, 2012 all of which are herebyincorporated by reference in their entirety.

TECHNICAL FIELD

The present disclosure relates to pressure sensors. More specifically,the present disclosure relates to miniaturized wireless pressuremonitoring devices, and related components and methods.

BACKGROUND

Biomedical engineering with real-time biological information, such aseye pressure, blood pressure, core body temperature and neural signals,has become useful in the research for identifying genetic variationsusceptibility to diseases. Animal-based research result is expected tomake a helpful impact in developing new treatment methods for similarhuman diseases. A miniature and implantable bio-sensing microsystem withwireless telemetry and RF powering is highly desirable to captureaccurate bio-signals and information. These CMOS-based devices haveadditional aspects for managing a large number of channels: signalmultiplexing, amplification, A/D conversion and filtering are doneon-chip. Developing the micro-system includes sensing methods study forvarious vital signals. Among the biological signals, eye pressure forGlaucoma patients is one of the more useful vital signals.

Glaucoma, the second leading cause of blindness, is a debilitatingdisease that affects millions of people. Current data suggests that 60million people have glaucoma, and that number is to reach 79 million bythe year 2020. Several risk factors put you at a disadvantage tocontract the disease, including ethnicity, age, and family history. Inglaucoma, death of optic nerve cells leads to blindness. Currently,doctors have very few options in detecting glaucoma, and many of themonly detect glaucoma after enough damage has occurred and it is too lateto prevent the disease. These include visual field testing, OcularCoherence Tomography, and Intra Ocular Pressure (IOP) monitoring.

The first two determine glaucoma after the damage has been completed,and can be used as an aid to monitor damage that has already beencompleted. IOP monitoring is the only method in which doctors canprevent the glaucoma from beginning and is the strongest knowncontributing factor to optic nerve cell death. Although it is clear thathigh IOP is harmful to the optic nerve, the exact relationships betweenthe kinetics, duration, and magnitude of IOP elevation and optic nervedamage are not known.

Doctors use several methods to examine IOP's glaucomatous effects.First, Goldmann tonometry (the gold standard) is a noninvasive method tomeasure the pressure inside the eye. Goldmann tonometry uses theImbert-Fick law that equates pressure based on flattening of the cornea.This leads to misdiagnosis from individual to individual based onphysiological characteristics. Studies conducted by Whitacre et al.found several errors occurring from using these indentation, ormathematical computations, to determine IOP. The only viable invasivemeasurement technique currently available is microneedle cannulation.This process involves inserting a needle, connected to a pressure gauge,into the anterior chamber of the eye and monitoring. This is a veryinvasive method of monitoring eye pressure, and if a physician needscontinuous data, it is infeasible. As IOP changes considerably evenwithin a single day, typical annual or semi-annual measurements at adoctor's visit can be misleading or uninformative.

The mouse is a well-suited mammalian model for deciphering themechanisms of this complex disease due to the similarity of itswell-known genome to the human. N-ethyl-N-nitrosourea (ENU) mutagenesisfollowed by phenotypic screening is proving to be a fruitful approach tostudying glaucoma. In this approach, there is considerable variabilityin age of onset of IOP elevation, even in mice with the same mutation.Thus, success is limited where IOP measurement data over many monthscannot be continuously and remotely assessed. In order to study themechanism of the disease, continuous monitoring of intraocular pressure(IOP) using an implanted sensor is useful. The diameter of the anteriorchamber in a mouse eye is approximately 3 mm and its depth is 75 micronsat the edge and 300 microns in the center. However, current IOP sensorsare not suitable to be used as a mouse eye implant due to their size.Furthermore, sensors that use near field inductive coupling do notprovide the sensitivity and the sensing distance needed for IOPmonitoring with a resolution of 1 mmHg from mice which are awake androaming within a cage. Therefore, a fundamental challenge lies in thecreation of a suitably sensitive sensor that will fit in the small spaceavailable in the eye of the mouse and provides the needed continuousmonitoring and sensing properties for use with mice in typicallaboratory cages.

In addition, implantable pressure sensors are useful with regards toother conditions. As one example, the breast implant market is 1-2billion dollars a year. Half of those sales are for silicone gelimplants. The primary concern is that when they rupture, there is no wayto detect the rupture without an MRI scan. The FDA recommends all womenreceive an MRI every 2-3 years at a cost of $1,500-$2,000 each.

As implantable devices become “smart”, using Application SpecificIntegrated Circuits (ASICs) and other electrical components to sensetheir surroundings, power consumption becomes an aspect of running thesedevices efficiently. Currently, there are several methods in which apower source is supplied to implantable devices. One includes the use ofhigh mAh batteries that allow the device to be powered for many years.Second, such as with implantable hearing aids, inductive coupling isused with an external power source coupling energy to the implanteddevice to stimulate the hair follicles inside the cochlea.

These two powering techniques, although acceptable, have drawbacks.Using a battery is efficient and allows for a long implant life, butonce that battery has died a second usually major surgery is required tooutfit a patient with a new battery. In the case of a cardiac pacemaker,this is costly and life threatening. Using inductive coupling, a staplein cochlear implants, is efficient in close range to power the device.However, there lies its drawback, that the external device should alwaysbe closely located and arranged for proper alignment and powering.

Improvements and alternatives are therefore needed in this field.

SUMMARY

According to one aspect, the present disclosure relates to a capacitivepressure sensor for monitoring fluid pressure within a patient isdisclosed, the pressure sensor comprising: a flexible substrateconfigured to bend so as to conform to a curved surface; a firstelectrode in abutting contact with the substrate, wherein the firstelectrode is configured to bend in conformity with the substrate; amembrane spaced from the first electrode, wherein the membranecomprises: a second electrode configured to be displaced toward thefirst electrode; and at least one flexible layer covering at least aportion of the second electrode; and a dielectric region between thefirst electrode and the membrane, wherein the pressure sensor has asensitivity of no less than about 0.3 fF/mmHg whether the substrate isin a flat orientation or a curved orientation.

According to another aspect, the present disclosure relates to acapacitive pressure sensor for monitoring fluid pressure within apatient, the pressure sensor comprising: a substrate; a first electrodein abutting contact with the substrate; a membrane spaced from the firstelectrode, wherein the membrane comprises: a second electrode configuredto bend toward the first electrode; and at least one flexible layercovering at least a portion of the second electrode; and a dielectricregion between the first electrode and the membrane having a depth of nogreater than about 7 microns when the membrane is in an uncompressedposition; wherein the pressure sensor has a sensitivity of no less thanabout 0.3 fF/mmHg. According to other aspects, the second electrode maybe thinner than the at least one flexible layer so as to readily bendwhen the at least one flexible layer is bent toward the substrate.According to other aspects, the second electrode may be sandwichedbetween two flexible layers. According to other aspects, the substrateand the at least one flexible layer may be polymeric. According to otheraspects, the substrate may comprise liquid crystal polymer and the atleast one flexible layer may comprise parylene. According to otheraspects, the sensitivity may be substantially constant over a range offrom about 0 mmHg to about 50 mmHg above atmospheric pressure. Accordingto other aspects, the perimeter of the substrate may define an area ofno greater than about 2 millimeters². According to other aspects, apressure sensing area of the sensor may be no greater than about 0.5millimeters². According to other aspects, the an additional flexiblelayer may cover at least a portion of the first electrode, wherein theadditional flexible layer and the at least one flexible layer of themembrane comprise the same material. According to other aspects, thedielectric region may comprise a gas. According to other aspects, thedielectric region may comprise air.

According to another aspect, the present disclosure relates to a systemfor monitoring fluid pressure within a patient, the system comprising: aflexible substrate configured to bend so as to conform to a curvedsurface, wherein the flexible substrate defines a maximum width of nogreater than about 0.7 millimeters; a capacitive pressure sensor coupledwith the substrate, wherein a maximum width of the pressure sensor is nogreater than the maximum width of the substrate, the pressure sensorcomprising: a first electrode in abutting contact with the substrate,wherein the first electrode is configured to bend in conformity with thesubstrate; a membrane spaced from the first electrode, wherein themembrane comprises: a second electrode configured to be displaced towardthe first electrode; and at least one flexible layer covering at least aportion of the second electrode; and a dielectric region between thefirst electrode and the membrane; and an integrated circuit coupled withthe substrate and electrically coupled with the pressure sensor, whereina maximum width of the integrated circuit is no greater than the maximumwidth of the substrate.

According to another aspect, the present disclosure relates to a methodof monitoring intraocular pressure in a mouse, the method comprising:implanting a pressure-sensing system within the anterior chamber of aneye of a mouse, wherein the pressure-sensing system comprises a flexiblesubstrate having an area of no greater than about 2 millimeters²; andreceiving data from the pressure-sensing system via equipment that isexternal to the mouse. According to other aspects the pressure-sensingsystem comprises a pressure-sensing area of no greater than about 0.5millimeters². According other aspects the pressure-sensing systemcomprises a capacitive pressure sensor that is integrated with theflexible substrate and comprises a flexible membrane, wherein theflexible substrate comprises liquid crystal polymer and the flexiblemembrane comprises parylene.

According to another aspect, the present disclosure relates to a systemfor monitoring fluid pressure within an eye, the system comprising: asubstrate defining a maximum width of no greater than about 0.7millimeters; a pressure sensor coupled with the substrate, wherein amaximum width of the pressure sensor is no greater than the maximumwidth of the substrate; an integrated circuit coupled with the substrateand electrically coupled with the pressure sensor, wherein a maximumwidth of the integrated circuit is no greater than the maximum width ofthe substrate; an antenna coupled with the substrate and electricallycoupled with the integrated circuit, the antenna and integrated circuitadapted to receive a first wireless signal for powering the system andtransmit a second wireless signal, the second wireless signal providingan indication of the fluid pressure; an LED coupled with the substrateand electrically coupled with the integrated circuit, wherein a maximumwidth of the LED is not greater than the maximum width of the substrate,wherein the light intensity of the LED output provides an indication ofpower being received by the system from a wireless external powersource. In certain embodiments, the antenna transmits the second signalat a harmonic of the frequency of the first signal, such as the thirdharmonic.

According to another aspect, the present disclosure relates to a systemfor monitoring fluid pressure within an eye, the system comprising: asubstrate defining a maximum width of no greater than about 0.7millimeters; a pressure sensor coupled with the substrate, wherein amaximum width of the pressure sensor is no greater than the maximumwidth of the substrate; an integrated circuit coupled with the substrateand electrically coupled with the pressure sensor, wherein a maximumwidth of the integrated circuit is no greater than the maximum width ofthe substrate; an antenna coupled with the substrate and electricallycoupled with the integrated circuit, the antenna and integrated circuitadapted to receive a first wireless signal for powering the system andtransmit a second wireless signal, the second wireless signal providingan indication of the fluid pressure, wherein the second wireless signalis transmitted at a harmonic of the first wireless signal, such as thethird harmonic.

According to another aspect, the present disclosure relates to a methodof forming a capacitive pressure sensor, the method comprising:providing a flexible polymeric substrate; depositing a first layer ofmetal on the polymeric substrate so as to form a first electrode; andforming a flexible membrane over the first layer of metal, wherein saidforming comprises: positioning a first polymeric layer above at least aportion of the first layer of metal; depositing a second layer of metalover the first polymeric layer; and depositing a second polymeric layerover at least a portion of the second layer of metal. The method mayfurther comprise depositing a layer of sacrificial photoresist over atleast a portion of the first layer of metal, wherein said positioning afirst polymeric layer above at least a portion of the first layer ofmetal comprises depositing the first polymeric layer over the layer ofsacrificial photoresist. The method may further comprise: etching anopening in the flexible polymeric substrate; and removing thesacrificial photoresist via the opening so as to provide a gap betweenthe first and second layers of metal in the region previously occupiedby the sacrificial photoresist. The method may further comprise4 sealingthe opening so as to enclose the gap. The method may further comprisedepositing a base layer of polymeric material over the first metallayer, wherein said positioning a first polymeric layer above at least aportion of the first layer of metal comprises bonding the firstpolymeric layer to the base layer of polymeric material. According tocertain aspects, the flexible polymeric substrate may comprise liquidcrystal polymer and each of the first and second polymeric layers maycomprise parylene.

According to another aspect, the present disclosure relates to a methodfor obtaining a measurement of pressure in an environment within asubject, comprising: irradiating an implanted pressure sensor system inthe subject by wirelessly transmitting radiowaves to the implantedpressure sensor system to power the system, wherein the implantedpressure sensor system comprises a pressure sensor, an integratedcircuit electrically coupled to the pressure sensor, and an antennaelectrically coupled to the integrated circuit; and receiving data fromthe implanted pressure sensor system by radiowaves, wherein the datarepresents information indicative of the pressure. The pressure sensorsystem may include a pressure sensor according to any embodiment orformed according to any method disclosed herein. According to certainaspects, the environment may be an eye.

According to another aspect, the present disclosure relates to a methodfor assessing the effects of an agent on intraocular pressure in an eyeof a subject, comprising: administering the agent to the subject;powering an implanted pressure sensor system in the eye by wirelesslytransmitting radiowaves to the implanted pressure sensor system, whereinthe implanted pressure sensor system comprises a pressure sensor, anintegrated circuit electrically coupled to the pressure sensor, and anantenna electrically coupled to the integrated circuit; and receivingdata from the implanted pressure sensor system by radiowaves, whereinthe data represents information indicative of the pressure. In certainembodiments, the agent is a biologically active molecule. In certainembodiments, the method for assessment may include comparing the datareceived from a control subject, wherein the control subject is the samesubject prior to administration of the agent or a second subject whichhas not been administered the agent.

According to another aspect, the present disclosure relates to anon-human animal having implanted therein a pressure sensor according toany embodiment disclosed herein.

BRIEF DESCRIPTION OF THE DRAWINGS

The present disclosure describes illustrative embodiments that arenon-limiting and non-exhaustive. Reference is made to certain of suchillustrative embodiments that are depicted in the figures, in which:

FIG. 1A is a top plan view of an embodiment of an implantable pressuresensing system;

FIG. 1B is a side elevation view of the system of FIG. 1A;

FIG. 10 is a cross-sectional view of a portion of the system of FIG. 1Ataken along the view line 1C-1C in FIG. 1A;

FIG. 2 is a cross-sectional view of a mouse eyeball;

FIG. 3 is an ultrasound image of a mouse eyeball that has been annotatedto depict various dimensions of the anterior chamber thereof;

FIG. 4 is a schematic cross-sectional view of an embodiment of apressure sensing system, such as that of FIG. 1A, implanted within theanterior chamber of a mouse eyeball;

FIG. 5A is a cross-sectional view of an embodiment of a pressure sensorin an uncompressed state;

FIG. 5B is a cross-sectional view of the pressure sensor of FIG. 5A in acompressed state;

FIGS. 6A-6I are schematic views depicting various stages of a method forfabricating an embodiment of a pressure sensor;

FIG. 7 is a top plan view of an embodiment of a pressure sensor that hasbeen fabricated using the method of FIGS. 6A-6I;

FIG. 8 is a cross-sectional view of the pressure sensor of FIG. 7;

FIG. 9 is a perspective view of an embodiment of a pressure controlledchamber that is used to test the pressure sensor of FIG. 7, wherein theinset image in the upper right corner depicts a close-up view of thepressure sensor within the chamber;

FIG. 10 is a plot of the capacitance of the pressure sensor of FIG. 7 asa function of pressure;

FIG. 11 is a plot of the change in capacitance of the pressure sensor ofFIG. 7 as a function of pressure, as compared with two theoreticalmodels for the same assuming a membrane residual stress of 0 MPa and 35MPa;

FIGS. 12A-12F are schematic views depicting various stages of anothermethod for fabricating an embodiment of a pressure sensor;

FIG. 13 is a 3-D image of an embodiment of a pressure sensor that hasbeen fabricated using the method of FIGS. 12A-12F;

FIG. 14 is a surface profile of the pressure sensor of FIG. 13;

FIG. 15 is a plot of the capacitance of the pressure sensor of FIG. 13as a function of pressure;

FIG. 16 is a plot of the change in capacitance of the pressure sensor ofFIG. 13 as a function of pressure, as compared with two theoreticalmodels for the same assuming a membrane residual stress of 0 MPa and 35MPa;

FIG. 17 is a schematic cross-sectional view of an embodiment of apressure sensing system implanted within the anterior chamber of a mouseeyeball, which further illustrates a radius of curvature of a surface ofthe anterior chamber and a theorized radius of curvature of a liquidcrystal polymer substrate positioned at that surface;

FIG. 18 is a perspective view of a pressure sensor having a flexiblesubstrate that is coupled to a mold having a curved surface, wherein thesubstrate is curved to conform to the curved surface;

FIG. 19 is a schematic cross-sectional view of the setup of FIG. 18taken along the view line 19-19; and

FIG. 20 is a plot comparing the performance of a pressure sensor whenits substrate is flat with its performance when positioned in a curvedsetup such as that of FIGS. 18 and 19.

FIG. 21A and 21B is schematic view of a third-order harmonic passive tagfor IOP monitoring (top view (left) and cross-sectional view (right))according to one embodiment.

FIG. 22A-H are schematic views depicting various stages of anothermethod for fabricating an embodiment of a compact-size tag formonitoring IOP inside the mouse eye according to one embodiment.

FIG. 23 illustrates a packaged parylene tag (left) and its implantationinside the mouse eye (right) according to one embodiment.

FIG. 24 is a schematic view of a setup for in-vivo testing according toone embodiment.

FIG. 25 is a plot of measured resonance frequency shift for the changeof IOP. (Normalized and fitted from the originally corrected data)according to one embodiment.

FIG. 26 is a block diagram of fully wireless implantable micro-systemaccording to one embodiment.

FIG. 27 is a top level schematic of a SAR ADC according to oneembodiment.

FIG. 28 is a top level schematic of comparator according to oneembodiment.

FIG. 29 is a preamplifier schematic design according to one embodiment.

FIG. 30 shows an auto zeroing clock generation circuit according to oneembodiment.

FIG. 31 is a schematic of a latch according to one embodiment.

FIG. 32 is a top level schematic of 10 bit SAR Logic according to oneembodiment.

FIG. 33 is a top level simulation result for 10 bit SAR ADC (input=400mV peak sinusoidal of 500 Hz) according to one embodiment.

FIG. 34 is a SAR ADC top level layout (320 um×240 um) according to oneembodiment.

FIG. 35 shows a LCP design for implantable IOP device according to oneembodiment.

FIG. 36 is a perspective view of a capsular tension ring IOP devicedesign according to one embodiment.

FIG. 37 is a base station functional block diagram according to oneembodiment.

FIG. 38 shows stained micrographs under 50× magnification used inhistological studies of LCP (a) after 7 days, (b) after 14 days, and (c)after 28 days; and parylene coated LTCC (d) after 7 days, (3) after 14days, and (f) after 28 days.

FIG. 39 illustrates measurements of fibrous encapsulation after 7 daysof implantation for various materials.

FIG. 40 illustrates measurements of fibrous encapsulation after 14 daysfor the materials of FIG. 39.

FIG. 41 illustrates measurements of fibrous encapsulation after 28 daysfor the materials of FIG. 39.

FIG. 42 shows a first surgical implantation of a tadpole IOP device.

FIG. 43 shows a first implantation of CTR IOP device.

FIG. 44 shows a dummy device implantation.

FIG. 45 shows a LED device active in mouse surgery.

FIG. 46 is a plot showing power received using RF powering for both LTCCand Silicon devices.

FIG. 47 shows a mouse-sized IOP device with surface mounted passivecomponents according to one embodiment.

FIG. 48 is a block diagram of a full system device according to oneembodiment.

FIG. 49 shows plots of a reflection coefficient characterization ofcomponents of obtaining and harvesting RF energy (a) real and (b)imaginary, according to one embodiment.

FIG. 50 shows a radio frequency powering setup to test devices in-vivoand ex-vivo according to one embodiment.

FIG. 51 shows a tadpole device based on a human IOP design.

FIG. 52 shows a large LED ring device of FR-4 according to oneembodiment.

FIG. 53 shows a small LED ring device of FR-4 according to oneembodiment.

FIG. 54 shows a further optimization onto LCP of an LED device accordingto one embodiment.

FIG. 56 shows a device size comparison to an “O” of a penny according toone embodiment.

FIG. 57 shows an IOP device above antenna according to one embodiment.

FIG. 55 shows an LED device wirelessly powered according to oneembodiment.

FIG. 58 shows a dummy sample being compressed for implantation accordingto one embodiment.

FIG. 60 is a plot showing the intensity area of a first LED device whentested ex-vivo according to one embodiment.

FIG. 61 is a plot showing the LED power consumption when tested ex-vivousing wireless RF powering for the device of FIG. 60.

FIG. 62 is a plot showing the intensity area of a second LED device whentested ex-vivo according to one embodiment.

FIG. 63 is a plot showing the LED power consumption when tested ex-vivousing wireless RF powering for the device of FIG. 62.

FIG. 59 shows a mouse animal model implanted with an LED deviceaccording to one embodiment.

FIG. 64 is a plot showing the intensity area of the LED device of FIG.60 when implanted in the anterior chamber of a mouse eye for 1 month.

FIG. 65 is a plot showing the power consumption of the LED device ofFIG. 60 when implanted in the anterior chamber of a mouse eye for 1month.

FIG. 66 shows the vascularization of a mouse eye following implantationof an LED device. Blue indicates vascularization. Purple is reflectance.

FIG. 67 shows a histological response 2 weeks post-mortem. Blue is thevascularization. Purple is reflectance.

FIG. 68 shows a 3D recreation of a mouse eye 2 weeks after implantationof an LED device. Color incorporates depth. Red constitutes tope whiteblue is farther away from lens.

FIG. 69 is a plot showing an analysis of efficiency of specific antennatypes connected to an ASIC rectifier circuit under different loadconditions according to one embodiment.

FIG. 70 is a schematic diagram show a condition for matching two compleximpedances that lie outside a 1+ω circle. Z_(L) and Z_(S) are loadimpedance and source impedance respectively. Z₁ and Z₂ are inductors orcapacitors for matching purposes.

FIG. 71 is a plot showing a measurement circuit output and arelationship between capacitive load and frequency output according toone embodiment.

FIG. 72 is a plot showing a comparison of measurement output andtransmission demodulation for a given capacitance according to oneembodiment.

FIG. 73 shows plots of output data of a measurement system using a MEMSpressure sensor in a pressure chamber according to one embodiment.

FIG. 74 is a block diagram of a coherent double sampling capacitance tovoltage converter according to one embodiment.

DETAILED DESCRIPTION

For the purposes of promoting an understanding of the principles of theinvention, reference will now be made to the embodiments illustrated inthe drawings and specific language will be used to describe the same. Itwill nevertheless be understood that no limitation of the scope of theinvention is thereby intended, such alterations and furthermodifications in the illustrated device, and such further applicationsof the principles of the invention as illustrated therein beingcontemplated as would normally occur to one skilled in the art to whichthe invention relates. At least one embodiment of the present inventionwill be described and shown, and this application may show and/ordescribe other embodiments of the present invention. It is understoodthat any reference to “the invention” is a reference to an embodiment ofa family of inventions, with no single embodiment including anapparatus, process, or composition that should be included in allembodiments, unless otherwise stated. Further, although there may bediscussion with regards to “advantages” provided by some embodiments ofthe present invention, it is understood that yet other embodiments maynot include those same advantages, or may include yet differentadvantages. Any advantages described herein are not to be construed aslimiting to any of the claims.

Various embodiments of pressure-sensing devices and systems aredisclosed herein, as are various components useful in the same andvarious methods of making and using the same. In certain embodiments,capacitive pressure sensors are significantly smaller than prior artdevices, yet are as sensitive as, or more sensitive than, those priorart devices. The capacitive pressure sensors thus can be deployed in agreater variety of destinations, as they can fit in more size-restrictedregions. In other or further embodiments, the capacitive pressuresensors are extremely thin and/or include a flexible substrate, such as,for example, a polymer-based material. The pressure sensors thus canconform to a surface curvature or profile at an implantation region. Forexample, a pressure sensor may be implanted within a cavity or chamberof an animal subject or, more particularly, a human or other mammaliansubject, such as, for example, within the anterior chamber of the eye,e.g. of a mouse. It shall be understood that the disclosed pressuresensors may also be implanted or otherwise operated in non-biologicalsmall scale environments where space or access is limited. The substratecan conform to at least a portion of a wall of the anterior chamber,which allows the sensor to more readily fit within the small space ofthe anterior chamber and/or which reduces or eliminates irritation,inflammation, and/or other complications that could result from aninflexible or rigid substrate body. In still other or furtherembodiments, the capacitive pressure sensors are biocompatible, suchthat they can be implanted within a subject without being rejected bythe body of the subject. As used herein, the term “subject” applies toany suitable animal (e.g., mammal) within which a pressure sensor is orcan be implanted. One or more of the foregoing advantages, and/or or oneor more other advantages, of various embodiments of pressure sensors andsystems will be apparent from the following discussion. Similarly,advantages of methods for making and using such pressure sensors willalso be evident.

Much of the following disclosure is directed to embodiments ofcapacitive pressure sensors that can be implanted in the anteriorchamber of the eye of a small mammalian subject, such as a mouse (see,e.g., FIGS. 2-4). In some embodiments, the capacitive pressure sensor issufficiently small for implantation in the eye of a mouse, sufficientlysensitive to obtain the desired pressure readings, and/or is capable oflong-term implantation in the eye of a mouse with little or noinflammation of the eye. Furthermore, certain embodiments of thepressure-sensing implants are configured for automated, remote, and/orcontinuous monitoring of the intraocular pressure (IOP) within the eyeof a subject, in certain embodiments a mouse eye or a human eye.

Some embodiments of the present disclosure pertain to methods formonitoring the response of a non-human animal. In some embodiments, themethod includes preparing an incision in the eye of the animal, andinserting an active pressure sensor through the incision into a spacewithin the eye. Preferably, the sensor is powered by electromagneticradiation, especially radio waves. Further, it is preferable that thesensor transmits pressure data by electromagnetic radiation, especiallyradio waves.

In yet other embodiments, the method includes providing a radio waveantenna on the sensor for the collection and transmission of radiowaves. Preferably, the antenna is collapsible to a small size compatiblewith the size of the incision. In some embodiments, the antenna isfabricated from a material that expands back to a predetermined shapethat is adapted and configured for collection and transmission of radiowaves within the biological space of the animal.

Yet other embodiments of the present disclosure pertain to a sensorassembly implantable with in a media, and useful for measuring thepressure within the media. In some embodiments, the sensor assemblyincludes a sensor providing a signal corresponding to the pressure themedia, a circuit for receiving the signal and providing an output to anantenna, the antenna being adapted and configured to not have a freeend. In those applications in which the media is within the biologicalspace of an animal, the lack of a free end makes the antenna morebiocompatible, since there is no free end to cause irritation.

Yet other embodiments the present disclosure pertain to a sensorimplantable in media within the biologic space of an animal. The antennais preferably fabricated from a shape memory material, such that it canbe collapsed to a first, smaller size prior to placement within thebiologic space, and then expand according to the shape memorycharacteristics to a second, larger shape that is more compatible withthe biologic space, and which in some embodiments is further adapted forthe transmission or receipt of radio waves within a predeterminedfrequency band.

Still further embodiments of the present disclosure pertain to anapparatus for measuring pressure within the media. In some embodimentsthe sensor is of the capacitance type, and is accompanied by a circuitthat provides for signal processing of the capacitance signal. In oneembodiment, the sensor interfaces with circuitry that provides thesensor signal at a higher order harmonic of the radiowave frequency thatwas utilized to power the sensor. In still further embodiments, thecircuitry includes a low-power, 10 bit analog to digital converter thatprovides a high resolution signal of the changes in capacitance.

Yet other embodiments of the present disclosure pertain to a method ofmeasuring a changing capacitance by use of a correlated double samplingcapacitance to voltage converter. In some embodiments, this signalprocessing can be modified externally on a sensor that has been placedwithin the media whose pressure is being measured. For example, in thoseapplications in which the base capacitance of the sensor varies, thecircuitry can include a processor and repeatedly programmable memorythat receives data corresponding to one or more circuit parameters.Receipt of this data can be used to vary these parameters duringoperation of the sensor. For example, the data can refer to parametersthat will affect the sensitivity of the circuit to 1/F noise, datarefers to the tuning of a reference capacitor, or data useful inmodifying characteristics of a common mode rejection circuit.

Various embodiments of the present disclosure demonstrate thefeasibility of IOP monitoring inside a mouse eye using a miniature 3rdorder harmonic tag or chip that is useful for large population glaucomastudies in the future. A novel packaging approach is introduced toimplement an ultra-small form factor implant that is thin enough to beimplanted inside a mouse eye. A microelectricalmechanical system (MEMS)capacitive pressure sensor is successfully integrated with aself-expandable Nitinol antenna and a light emitting diode (LED) withinthe small tag. Other embodiments include use of the miniature 3^(rd)order harmonic tag in an eye of a human or other animal.

Although this disclosure discusses this illustrative context in detail,it should be appreciated that the pressure sensors are readily used inother contexts. For example, as discussed further below, certainpressure sensors may be used within the heart of a subject (e.g., ahuman) and/or within the vasculature of the subject. Some pressuresensors can be used to monitor the pressure of cerebrospinal fluid.Others can be positioned within silicone breast implants to sense achange in pressure due to rupture. A wide variety of other applicationsare also possible.

FIGS. 1A-1C depict one embodiment of a pressure-sensing implant, orpressure-sensing system 100, that is configured to be implanted in theeye of an animal. In particular, the pressure-sensing system 100 isconfigured to be implanted in the eye of a mouse. It shall be understoodthat the system 100 may also be sized and adapted for implantation in ahuman eye or other animal eye. The pressure-sensing system 100 includesa substrate 110 that has multiple components mounted thereto and/orintegrated therewith. In the illustrated embodiment, thepressure-sensing system 100 includes an integrated circuit (IC) 112, apressure sensor 114, and an antenna 116 that are mounted to orintegrally formed with the substrate 110. In view of the presentillustrative context in which the pressure-sensing system 100 is used,the pressure sensor 114 may also be referred to as an IOP monitoringsensor.

As shown in FIG. 1C, the integrated circuit 112 can be electricallycoupled with each of the pressure sensor 114 and the antenna 116 so asto be able to electrically communicate therewith. In particular,separate portions of the pressure sensor 114 can be coupled with theintegrated circuit 112 via a pair of electrical leads 120, 122, asdiscussed further below. The pressure sensor 114 can be configured toprovide data (e.g., capacitance values) to the integrated circuit 112,which can be configured to store the data, derive information from thedata (e.g., derive calculated or estimated pressure values from thecapacitance values), and/or deliver the data and/or the derivedinformation to the antenna 116 for delivery to a remote data storageand/or analysis system (e.g., a computer or other electronic device, notshown).

In the illustrated embodiment, the integrated circuit 112 interfaceswith the antenna 116, which in one embodiment can be an inductive loopantenna, via a metal-insulator-metal (MIM) capacitor 124, the capacitor124 providing for radio-frequency (RF) matching between the antenna 116and the integrated circuit 112. A first electrode, or capacitive plate126, of the MIM capacitor 124 is coupled with the integrated circuit 112via at least one electrical lead 127, and second electrode, orcapacitive plate 128, of the MIM capacitor 124 is coupled with theantenna 116 via at least one additional electrical lead 129. Otherand/or additional suitable electrical connections between the integratedcircuit 112 and the antenna 116 are possible.

For a desired RF power transfer between two components (e.g., maximum oroptimized power transfer), a conjugate matching may be used. Connectingthe MIM capacitor 124 to the highly inductive loop antenna 116 canprovide approximate conjugate matching between the antenna and theintegrated circuit, which can provide a desirable power transfer (e.g.,can maximize or optimize power transfer) from the antenna to the RFmatching circuit in the integrated circuit 112. In certain embodiments,there may be no other electrical components between the antenna 116 andthe integrated circuit 112, and the RF powering circuit can be embeddedin the integrated circuit 112. In certain embodiments, the diameter ofthe loop antenna 116 may be sized to circumscribe the iris or cornea. Incertain embodiments, the values for the matching may be determined asshown in Appendix A.

In various embodiments, the antenna 116 is capable of transmittingsignals provided by the integrated circuit 112 to any suitable receiver(not shown) that is remote from the system 100 (e.g., at a positionexternal to the subject within which the system 100 has been implanted).The signals may then be stored and/or processed remotely by any suitableprocessor (not shown). In such a manner, data regarding IOP within theeye of a mouse or other subject can be monitored on a regular and/orcontinuous basis. Such monitoring may be automated. In other or furtherembodiments, the antenna 116 is capable of receiving signals from asource (e.g., a computer or other electronic device) that is remote fromthe system 100. For example, the system 100 may be remotely programmableand/or controllable by the external source. In still other or furtherembodiments, the antenna 116 may be used to provide power to the system100. Stated otherwise, the antenna 116 can be used to relay informationto and/or from the integrated circuit 112, and in other or furtherembodiments, the antenna 116 can be used to provide power to theintegrated circuit 112. In some embodiments, RF power received by theantenna can be transferred to an RF powering circuit within theintegrated circuit 112. The RF powering circuit can be a rectifier,which can include diodes and capacitors. The rectifier can convert theRF power to a DC voltage that is used to power the integrated circuit112. In some embodiments, wireless powering and communication such asthat disclosed in Chow et al., “Fully Wireless ImplantableCardiovascular Pressure Monitor Integrated with a Medical Stent,” IEEETransactions on Biomedical Engineering, Vol. 57, No. 6, 1787-96 (June2010), the entire contents of which are hereby incorporated byreference, may be used.

In the illustrated embodiment, the antenna 116 comprises aself-expandable loop of shape-memory material, such as nickel titanium(Nitinol™). Use of a shape-memory material can aid with implantation ofthe system 100, as the antenna 116 may have a compressed configurationduring implantation, and after implantation, may naturally expand to itslooped shape, for example, due to the body temperature of the subject.The compressible and expandable nature of the antenna 116, and the smallsize of the substrate 110, can permit the system 100 to be easilyinjected into place. For example, in some embodiments, the system 100can be packaged in a compressed state within a microneedle, such as thatdisclosed in John et al., “Intraocular Pressure in Inbred MouseStrains,” Investigative Ophthalmology & Visual Science, Vol. 38, No. 1,249-53 (January 1997), the entire contents of which are herebyincorporated by reference herein. The system 100 can be urged from theinterior of the microneedle into the anterior chamber, and can then bepermitted to transition to an expanded state.

The integrated circuit 112 can comprise any suitable configuration orconstruction. The integrated circuit 112 can be an application-specificintegrated circuit (ASIC), which can be programmed or otherwiseconfigured to achieve any of the actions described herein, asappropriate. In various embodiments, the integrated circuit 112comprises a complementary metal-oxide-semiconductor (CMOS) integratedcircuit (IC). The integrated circuit 112 can be programmed for anysuitable functionality in any suitable manner. As shown in FIG. 1 C, theintegrated circuit 112 may be coupled with other portions of the system100 via any suitable adhesive 130. For example, in the illustratedembodiment, the adhesive 130 comprises a layer of Z-axis anisotropicconductive adhesive (ACA). The integrated circuit 112 may desirably berelatively thin so as to reduce the bulk or overall footprint of thesystem 100. The integrated circuit 112 and the adhesive 130 can define athickness T₂. In various embodiments, the thickness T₂ can be no greaterthan about 100, 125, 150, 175 or 200 microns, or can be within a rangeof from about 100 to about 200 microns or from about 125 to about 175microns. In some embodiments, which may be particularly well-suited foruse with mouse eyes, the thickness T₂ can be about 150 microns.

As shown in FIGS. 1A and 1B, the integrated circuit 112 may bepositioned at a generally central region of the expanded system 100. Forexample, in the illustrated embodiment, the integrated circuit 112 ispositioned at a longitudinal end of the elongated substrate 110 (seeFIG. 1A), or stated otherwise, is positioned closer to one longitudinalend of the substrate 110 than the other longitudinal end thereof. Theantenna 116 can be coupled with the system 100 at the oppositelongitudinal end of the substrate 110, and may encircle, circumscribe,or otherwise encompass the portion of the substrate 110 to which theintegrated circuit 112 is coupled. The integrated circuit 112 can berelatively thick, as compared with other components of the system 100(e.g., the pressure sensor 114 and the antenna 116). Accordingly, thesystem 100 can be generally thinner at its periphery and thicker at acentral region thereof. Such an arrangement can correspond to the shapeof the anterior chamber of a mouse eye within which the system 100 canbe implanted, as discussed further below.

Much of the following discussion focuses on the substrate 110 and thepressure sensor 114. The substrate 110 can desirably permit the system100 to bend, flex, bow, curve, or otherwise conform to a rounded orotherwise curved surface contour. The substrate 110 may also havemalleable, non-abrasive, or atraumatic edges and/or corners that arecapable of reducing or eliminating irritation or inflammation. Forexample, the substrate 110 can be thin and/or can comprise a flexiblematerial. The pressure sensor 114 can define a very small area, yet canbe highly sensitive, e.g., to fluid pressure fluctuations. These and/orother properties of the substrate 110 and the pressure sensor 114 can beparticularly useful in implanting the system 100 in the eye of a mouse,as discussed further below, although it will be understood that system100 can also be implemented in the eye of a human or other animal aswell.

The substrate 110 can be ultra-thin and flexible. In certainembodiments, the substrate 110 can comprise a single layer of a materialthat is biocompatible, chemically inert, flexible, conformable, strong,durable, and/or readily shaped via micro-machining. The substrate 110may have desirable electro-mechanical properties, such as a relativelylow dielectric constant, a low loss tangent, a low Young's modulus,and/or suitable compatibility with conventional micro-fabricationprocesses. For example, the substrate 110 can comprise a polymericmaterial, such as liquid crystal polymer (LCP).

A maximum thickness T₄ of the substrate 110 (see FIG. 1C) can beextremely small. In various embodiments, the thickness T₄ is no greaterthan about 15, 20, 25, 30, 35, 40, 45, or 50 microns, or is within arange of from about 15 to about 50 microns, from about 20 to about 30microns, or from about 24 to about 26 microns. In some embodiments, thethickness T₄ is about 25 microns, which can be desirable forapplications in which the system 100 is implanted in a mouse eye. Such athickness can permit at least a portion of the substrate 110 to readilyconform to a posterior wall of the anterior chamber of a mouse eye, yetprovide sufficient strength and support for other components of thesystem 100 that are mounted thereto and/or integrated therewith.Moreover, a relatively thick substrate 110 can be more resistant todeformation (e.g., localized deformation) than other portions of thesystem 100, which can aid in operation of the pressure sensor 114, asdiscussed further below.

With reference to FIG. 1A, the substrate 110 can define an extremelysmall area. In some embodiments, the substrate 110 is substantiallyrectangular, although other suitable shapes are possible. In variousembodiments, a length L₁ of the substrate 110 can be no greater thanabout 0.5, 0.75, 1.0, 1.1, 1.2, 1.3, 1.4, or 1.5 millimeters, or can bewithin a range of from about 0.5 to about 1.5, from about 0.75 to about1.5, or from about 1.0 to about 1.5 millimeters. In various embodiments,a width W₁ of the substrate can be no greater than about 0.25, 0.5, 0.6,0.65, 0.7, 0.75, 0.8, 0.9, or 1.0 millimeters, or can be within a rangeof from about 0.25 to 1.0, 0.5 to about 1.0, or 0.6 to about 0.8millimeters. In various embodiments, a perimeter of the substrate 110defines an area of no greater than about 0.1, 0.2, 0.3, 0.4, 0.5, 0.6,0.7, 0.8, 0.9, 1.0, 1.25, 1.5, or 2.0 millimeters². In some embodiments,which may be particularly well-suited for use with mouse eyes, thelength L₁ can be about 1.3 millimeters and the width W₁ can be about 0.7millimeters.

As previously discussed, the substrate 110 can be relatively flexible.For example, in various embodiments, the substrate 110 can have aYoung's modulus that is within a range of from about 10 to about 40 GPa,or that is no greater than about 10, 20, 30, or 40 GPa.

The pressure sensor 114 can comprise a microelectromechanical system(MEMS) capacitive pressure sensor. In certain embodiments, the pressuresensor 114 can be integral with the substrate 110, such that at least aportion thereof comprises a polymeric material. Moreover, as furtherdiscussed below, the pressure sensor 114 can comprise other polymericmaterials that differ from those used in the substrate 110. Accordingly,the pressure sensor 114 may be referred to as a polymer-based pressuresensor.

With reference to FIG. 1C, the pressure sensor 114 can include one ormore layers of a coating material, or a form factor material 140, thatis positioned above and/or is in abutting contact with the substrate110. The material 140 can have desirable form factor properties forforming a membrane 142, which is configured to be displaced relative tothe substrate 110, as discussed further below. The material 140 candesirably be chemically and/or biologically inert and biocompatible. Insome embodiments, the material 140 can be suitable for multi-layerprocessing (e.g., via any suitable deposition process, such as vacuumdeposition), and it can provide a conformal coating. The material 140may be deposited as one or more thin layers or films, and it may exhibita low coating temperature. In some embodiments, the material 140 may betransparent. The material 140 can desirably be flexible, or have arelatively low stiffness (e.g., compared with a stiffness of thesubstrate 110), which can permit the membrane 142 to bend, deform, orotherwise deflect toward the substrate 110 when pressure is applied tothe implanted system 100. As previously discussed, the substrate 110 maynot bend or deform due to the increased pressure, or may deform to amuch lesser extent, such that the substrate 110, and a lower electrodesupported thereby, can serve as a reference against which changes in themembrane 142 can be measured. In the illustrated embodiment, thematerial 140 comprises parylene. In various embodiments, the material140 has a Young's modulus of no greater than about 3, 3.5, 4, or 4.5GPa. In some embodiments, which may be particularly well-suited for usewith mouse eyes, the material 140 comprises parylene that has Young'smodulus of about 4 GPa. In various embodiments, a thickness T₃ of thematerial 140 can be no greater than about 1, 2, 3, 4, 5, 6, 7, 8, 9, or10 microns, or can be within a range of from about 1 to 10 microns. Inthe illustrated embodiment, the thickness T₃ is about 5 microns.

The pressure sensor 114 can further include a pair of electrodes orcapacitive layers or plates 144, 146 that are spaced or separated fromeach other. The plates 144, 146 can comprise any suitable conductivematerial, and at least the upper plate 144 may be extremely thin andflexible. In some embodiments, at least the upper plate 144 can comprisea thinly deposited layer of Ti/Au (titanium/gold). The plates 144, 146can be electrically coupled with the integrated circuit 112 via theleads 120, 122, respectively. In some embodiments, the upper plate 144and its lead 120 are sandwiched between layers of the material 140(e.g., parylene). The lower plate 146 and its lead 122 may also have alayer of the material 140 deposited thereon.

The pressure sensor 114 can define an insulating or dielectric region148 between the capacitive plates 144, 146. In the illustratedembodiment, the dielectric region 148 comprises an air-filled gap.However, it shall be understood that the dielectric region 148 maycomprise any material (e.g., gas, liquid, solid, gel) which allows theupper or lower plates 144, 146 to move or flex relative to one another,thereby creating a change in capacitance of the sensor 114. The specificmedium or material of the dielectric region 148 may be chosen based onthe needs of the particular application and dynamic range desired.Operation of the pressure sensor 114 and properties of the membrane 142and dielectric region 148 are discussed further below.

As previously discussed, embodiments of the system 100 can bewell-suited for implantation into the eye of a mouse for monitoring ofthe IOP within the eye. As shown in FIGS. 2-4, a mouse eyeball 200includes an anterior chamber 210 that includes a bowed or curvedposterior wall 212. The anterior chamber 200 is crescent-shaped incross-section, such that it defines a non-uniform thickness. A maximumthickness of the anterior chamber 200 is about 300 microns, whereas athickness near the edges thereof is about 75 microns. A maximum diameterof the anterior chamber 200 is about 3 millimeters. Embodiments of thesystem 100 are capable of fitting within the limited volume of theanterior chamber 200 and of conforming to the shape of the posteriorwall 212.

For example, with reference to FIGS. 1A, 1B, and 4, a maximum diameterof the system 100, which is represented by the length L₂, can be nogreater than about 2 millimeters. A maximum thickness T₁ of the system100 can be no greater than about 250 microns, with the maximum heightbeing at or near the center of the system 100. Accordingly, the system100 can fit well within the anterior chamber 200. In other embodiments,the length L₂ can be no greater than about 1, 1.5, 2.5, or 3millimeters, and in thickness T₁ can be no greater than about 100, 150,200, or 300 microns.

In certain embodiments, such as that schematically depicted in FIG. 4,the area available for the pressure sensor 114 is less than about 0.5millimeters² (e.g., fits within a square region measuring about 700microns by about 700 microns). Additionally, the pressure sensor 114 canhave a limited thickness, and may be no greater than about 150 microns.Further complicating the pressure sensing is the fact that the posteriorwall 212 is curved. It thus can be desirable for the substrate 110 tocurve so as to conform to the posterior wall 212 and/or to be flexibleor otherwise atraumatic so as to reduce, minimize, or eliminate damageto the eye tissue.

As previously mentioned, the maximum width W₁ of the substrate 110 canbe no greater than about 700 microns, in the illustrated embodiment, andthe antenna 116 can be originally packaged in a narrow or constrictedstate, and may self-expand after implantation. Accordingly, the system100 can be well suited for implantation through an incision that is nogreater than about 700 microns in length. This can minimize damage ofthe eye tissue during surgery and permit re-sealing of the incisionwithout sutures.

FIGS. 5A and 5B illustrates a portion of another embodiment of apressure-sensing implant or system 300 that includes a pressure sensor314, which can resemble the previously discussed pressure-sensing system100 and pressure sensor 114, respectively, in certain respects. Forexample, the pressure-sensing system 300 can closely resemble thepressure-sensing system 100, aside from any differences in theirrespective pressure sensors 314, 114. Accordingly, like features aredesignated with like reference numerals, with the leading digitsincremented to “3.” Relevant disclosure set forth above regardingsimilarly identified features thus may not be repeated hereafter.Moreover, specific features of the pressure-sensing system 300 or thepressure sensor 314 may not be shown or identified by a referencenumeral in the drawings or specifically discussed in the writtendescription that follows. However, such features may clearly be thesame, or substantially the same, as features depicted in otherembodiments and/or described with respect to such embodiments.Accordingly, the relevant descriptions of such features apply equally tothe features of the system 300 and the pressure sensor 314. Any suitablecombination of the features and variations of the same described abovewith respect to the system 100 and the pressure sensor 114 can beemployed with the system 300 and the pressure sensor 314, respectively,and vice versa. This pattern of disclosure applies equally to furtherembodiments depicted in subsequent figures and described hereafter.

The pressure sensor 314 comprises a capacitive pressure sensor that hasa high sensitivity to pressure changes, a low sensitivity to temperaturefluctuations, and a low power consumption. The pressure sensor 314 isformed atop a substrate 310, such as the substrate 110 discussed above.The pressure sensor 314 includes a diaphragm or membrane 342, whichincludes an upper electrode or capacitive plate 344 that is sandwichedbetween coating material or form factor material 340, such as thematerial 140 discussed above. The pressure sensor 314 further includes alower electrode or capacitive plate 346, which is in abutting contactwith the substrate 310. A dielectric region 348 is positioned betweenthe plates 344, 346. Again, it shall be understood that the dielectricregion 348 may comprise any material (e.g., gas, liquid, solid, gel)which allows the upper or lower plates 344, 346 to move or flex relativeto one another, thereby creating a change in capacitance of the sensor348. The specific medium or material of the dielectric region 348 may bechosen based on the needs of the particular application and dynamicrange desired. In the illustrated embodiment, the lower plate 346 doesnot include a coating material 340 over an upper surface thereof in thevicinity of the dielectric region 348.

The pressure sensor 314 is shown in a cross-sectional side elevationview. As viewed from above, the membrane 342 portion of the pressuresensor 314 is substantially circular. As the pressure of the environmentthat surrounds the pressure sensor 314 increases, the membrane 342 canbe deflected toward the substrate 310, as can be seen by comparing FIG.5A to 5B. This deflection changes the distance, and therefore thecapacitance, between the fixed bottom electrode 346 and the flexible topelectrode 344. The deflection of the circular membrane 342 and thechange in capacitance of the pressure sensor 314 can be estimated usingknown methods. For example, suitable methods for estimating the changein capacitance for a deflected circular membrane are disclosed inSchellin et al., “Measurement of the mechanical behavior ofmicromachined silicon and silicon-nitride membranes for microphones,pressure sensors and gas flow meters,” Sensors and Actuators A, vol. 41,287-92 (1994) and in Fragiacomo et al., “Analysis of small deflectiontouch mode behavior in capacitive pressure sensors,” Sensors andActuators A, vol. 161, pp 114-19 (2010), the entire contents of whichare incorporated by reference in their entirety herein. Generally, oneor more complicated or nonlinear formulas, relationships, or algorithmscan be used to relate a change in capacitance for a circular capacitorhaving known physical properties, such as the properties discussedbelow. While the embodiments discussed herein use circular geometriesfor the pressure sensor 314, it is noted that other geometries are alsopossible (e.g., square, rectangular, etc.).

The deflection caused by the applied pressure can be determined by therigidity and residual stress of the membrane 342. For example, themaximum deflection w, at the center of the membrane 342 can bedetermined as follows

$\begin{matrix}{w_{0} = \frac{P\; a^{4}}{{64D} + {4\; \sigma \; {ha}^{2}}}} & ( {{Equation}\mspace{14mu} 1} )\end{matrix}$

where P is the applied pressure to the membrane 342, D is the flexuralrigidity of the membrane 342, a is the radius of the upper electrode 344(which may be roughly the same as the radius of the membrane 342), h isthe thickness of the membrane 342, and a is the residual stress of themembrane 342. The flexural rigidity D can be expressed as

$\begin{matrix}{D = \frac{{Eh}^{3}}{12( {1 - v^{2}} )}} & ( {{Equation}\mspace{14mu} 2} )\end{matrix}$

where E is the Young's modulus of the membrane 342 and v is thePoisson's ratio of the membrane 342. The residual stress of the membrane342 can highly depend on the fabrication process conditions. Forexample, the residual stress D of thin parylene film that hasexperienced conventional micro-fabrication processes can be within arange of from about 30 to about 50 MPa. When considering the effect ofthe rigidity and residual stress on the membrane 344, the equation todescribe the deflection of the membrane is as follows:

$\begin{matrix}{{w(r)} = {w_{0}\lbrack {1 - ( \frac{r}{a} )^{2}} \rbrack}^{2}} & ( {{Equation}\mspace{14mu} 3} )\end{matrix}$

where r is the distance from the center of the membrane. The capacitanceat a given pressure is estimated from the deflection and the chargeintegral across the capacitor:

$\begin{matrix}{C = {2\; \pi \; ɛ{\int_{0}^{a}{\frac{r}{d - {w(r)}}{dr}}}}} & ( {{Equation}\mspace{14mu} 4} )\end{matrix}$

where d is the depth of the gap 348 between the top and bottom plates344, 346 in the absence of a net compressive pressure (e.g., in theuncompressed state shown in FIG. 5A).

As previously discussed, in some embodiments, the substrate 310 candesirably have a maximum width W₁ (see FIG. 1) that is no greater thanabout 700 microns, which can be desirable for insertion of thepressure-sensing implant into the eye of a mouse. Accordingly, in someembodiments, the radius of the membrane 342 is no greater than 350microns, such that the sensing area is no greater than about 0.38millimeters². More generally, in various embodiments, the sensing areacan be no greater than about 0.3, 0.4, or 0.5 millimeters².

Accordingly, fabrication methods for the pressure sensor 314 (and moregenerally, for the pressure-sensing device 300), and the dimensions andcomposition of the pressure sensor 314, can be selected, adjusted,determined, or otherwise result from a balancing of properties ofcomponents of the pressure sensor 314 and/or operational factors relatedthereto. For example, the depth d of the gap 348 and the thickness h ofthe membrane 342 can be determined from and/or limited based on theprocesses used in their fabrication and/or the desired sensitivity ofthe pressure sensor 314.

As a further example, it can be desirable to position the upper andlower electrodes 344, 346 as close as possible to each other so as tothereby improve sensitivity of the pressure sensor 314. However, in someinstances, it can be undesirable to position the electrodes 344, 346 soclose to each other such that, at higher pressures encountered withinthe eye of the mouse (or any other environment in which the sensor 314may be positioned), the membrane 342 contacts the lower electrode 344.Such contact can skew or alter the manner in which the capacitance ofthe electrodes 344, 346 varies with further increases in pressure, or itmay significantly curtail or even prevent any change in capacitance atthese increasingly higher pressures. Accordingly, the contact can eithercomplicate readings at the higher pressures or it may result in an upperlimit (e.g., “saturation level”) of pressure readings, thereby limitingthe overall sensitivity or dynamic range of the pressure sensor 314.

Contact between the membrane 342 and the lower electrode 346 at higherpressures could also cause the membrane 342 to undesirably adhere to thelower electrode 346. Accordingly, when a high pressure of theenvironment surrounding the pressure sensor 314 that caused the adhesionis subsequently reduced, the membrane 342 could undesirably remain stuckto the lower electrode 346, which would inhibit pressure readings at thelower pressures and could, in some instances, permanently impedeoperation of the pressure sensor 314 (e.g., should the membrane 342remain adhered to the lower electrode 346). Therefore, in manyinstances, it can be desirable for various parameters of the sensor 314,such as the depth d of the gap 348, the thickness h of the membrane 342,and/or the radius a of the upper electrode 346 of the membrane 342 to beselected so as to prevent contact between the membrane 342 and the lowerelectrode 346 at the highest pressures that are likely to be encounteredin the environment into which the pressure sensor 314 is implanted,while maintaining a desired sensitivity of the sensor 314. In thecontext of mouse eyes, the highest IOP likely to be encountered withinthe anterior chamber can be about 70 mmHg above atmospheric pressure,although even higher pressures may be encountered in some cases. In thecontext of a human eye, the highest IOP likely to be encountered isabout 80 mmHg. In the context of other environments (e.g., heart,vasculature, spinal fluid), higher or lower pressure ranges may beencountered. In various embodiments, the depth d of the gap 348 can beno greater than about 3, 4, 5, 6, or 7 microns, no less than about 3, 4,5, 6, or 7 microns, within a range of from about 3 to about 7 microns,within a range of from about 4 to about 6 microns, or can be about 5microns.

Discussed hereafter are illustrative methods for fabricating pressuresensors, such as the pressure sensors 114, 314 described above. Althoughthe methods are focused on the creation of pressure sensors, it isunderstood that the methods can be altered and/or extended to includefabrication of the pressure-sensing systems 100, 300 in theirentireties. For example, although procedural steps are not specificallydiscussed with respect to the coupling of integrated circuits orantennas with substrates, it is understood that such coupling ispossible at any suitable stage of the fabrication of the pressuresensors 114, 314.

Some of the illustrative methods discussed hereafter employ sacrificialphotoresist (PR) procedures, while others employ parylene membranetransfer procedures (e.g., parylene-to-parylene bonding). In either setof methods, a thin and biocompatible, conductive layer of Ti/Au issandwiched between parylene layers to form a membrane. Other or furtherconductive layers are possible (e.g., one or more other metal layers),but this could result in changes to other parameters of the pressuresensors (such as gap thickness, membrane thickness, etc.). Thesandwiched Ti/Au structure can be much more flexible than a metal orsilicon membrane of the same thickness. The thickness of the parylenelayers of the membrane can be selected or otherwise determined or formedso as to prevent the membrane from collapsing or tearing duringfabrication. The sandwiched Ti/Au structure can be extremely thin, whichcan yield a very low effective Young's modulus of the membrane, thusmaking the membrane easier to deflect (e.g., more sensitive to pressurefluctuations). In various embodiments, the capacitive pressure sensorscan be sufficiently sensitive to detect a pressure change of no lessthan about 0.5, 0.75, 1, 1.5, or 2 mmHg. For example, in someembodiments, the pressure sensors can desirably detect pressure changesof no less than about 1 mmHg. In various embodiments, the integratedcircuit that is used can also assist in the sensitivity of the pressuresensing system. Accordingly, it can be desirable to use a suitable ASICin connection with the physically sensitive pressure sensors, in someembodiments.

FIGS. 6A-6I depict an illustrative fabrication process that can be usedto create a capacitive pressure sensor 414 (FIG. 61), such as thecapacitive pressure sensors 114, 314 discussed above. Specificdimensions, materials, and other features are provided with respect tothis illustrative process. While such specific dimensions, materials,and features may constitute independently patentable subject matter, itshould be understood that other suitable dimensions, materials, andfeatures may be used in fabricating other capacitive pressure sensors inaccordance with other embodiments of the present disclosure, and thussuch dimensions, materials, and features are not necessarily meant to belimiting.

With reference to FIG. 6A, a photoresist (PR) layer 460 is applied to aflexible LCP substrate 410, which can have a thickness of 25 microns (orother suitable dimensions discussed above). The PR layer 460 cancomprise AZ1518, and can be spin-coated and hard-baked onto thesubstrate 410 for use as an adhesion layer. In certain embodiments, thesubstrate 410 can comprise UL TLARAM 3850, which is available fromRogers Corporation of Chandler, Ariz. The fabrication process for theLCP film can be similar to conventional Si-based MEMS processing. Forease of handling, the thin, flexible LCP substrate 410 and sacrificialPR layer 460 can be attached to a silicon carrier wafer 462.

One or more layers of Ti/Au may be used as metallic or conductive layersfor one or more of an upper and lower electrode 444, 446 (FIG. 61).Ti/Au may be selected due to, for example, its biocompatibility. Withreference to FIG. 6B, a layer of Ti/Au 464 can be sputtered onto (orotherwise applied to) a surface of the substrate 410 in an initial stageof formation of the lower electrode. To improve the Ti/Au adhesion, theLCP can be exposed to Ar (argon) ion bombardment in the sputteringsystem (which may be a Perkin Elmer 2400 sputter system) for 2 minutes.Ar ion bombardment under vacuum generates a TiC layer between Ti andLCP, which can play a role as a good adhesion layer. Then, the conductoror Ti/Au layer 464 is deposited. The Ti/Au layer 464 may have athickness of about 0.5 microns. The Ti/Au is patterned with wet etchingto form the lower electrode 446 and an associated electrical lead b 422.In the illustrated embodiment, a pathway, channel, or opening 466 isprovided in the Ti/Au layer 464, which is further discussed below.

With reference to FIG. 6C, a sacrificial PR layer 470 is deposited overat least a portion of the Ti/Au layer 464 using standard lithographyprocesses. The sacrificial PR layer 470 can define a thickness thatcorresponds to the desired thickness of a dielectric region or gap 448of the pressure sensor 414 (see FIG. 61). The specific medium ormaterial of the dielectric region 148 may be chosen based on the needsof the particular application and dynamic range desired. As discussedfurther below, the sacrificial PR layer 470 is eventually removed tothereby form the cavity or gap 448. In various embodiments, thesacrificial PR layer 470 can have a thickness of no greater than about1, 1.5, 2, 2.5, or 3 microns. In the illustrated embodiment, thethickness is about 2 microns. The sacrificial PR layer 470 can compriseany suitable PR material, such as, for example, MICROPOSIT SC1827, whichis available from Rohm and Haas Electronic Materials LLC of Marlborough,Mass.

As shown in FIG. 6D, a composite membrane 442 can then be created. Afirst parylene layer 472 can be deposited over the sacrificial PR layer470. The first parylene layer 472 can have any suitable thickness, suchas a thickness of no more than about 1, 2, 3, or 4 microns. In theillustrated embodiment, the first parylene layer 472 has a thickness ofabout 3 microns. The first parylene layer 472 can be deposited in anysuitable manner, such as via a PDS 2010 system, which is available fromSpecialty Coating Systems of Indianapolis, Ind. In some embodiments,before metal deposition onto the first parylene layer 472, the samplecan be etched with oxygen plasma (e.g., at 100 Watts for 20 seconds) ina Plasma-tech reactive-ion etching (RIE) system to improve Ti/Auadhesion on the parylene.

An upper Ti/Au layer 474 is deposited on the parylene layer 472 andpatterned so as to form the upper electrode 446 and an associatedelectrical lead 420. The electrode portion of the Ti/Au layer 474 can beof any suitable thickness, depending on the desired properties of thesandwiched membrane 442. For example, the Ti/Au layer 474 can have athickness of no more than about 0.2, 0.3, or 0.4 microns. In theillustrated embodiment, the Ti/Au layer 474 has a thickness of about 0.2microns.

A second parylene layer 476 can be deposited over at least a portion ofthe Ti/Au layer 474. The second parylene layer 476 can have any suitablethickness, such as a thickness of no more than about 1, 2, 3, or 4microns. In the illustrated embodiment, the second parylene layer 476has a thickness of about 2 microns.

With reference to FIG. 6E, carrier wafer 462 can be separated from anupper portion 480 of the assembly by soaking the assembly in acetone.Once the upper portion 480 of the assembly has been released from thecarrier wafer 462, it is flipped over and re-attached to the carrierwafer 462 via the spin-coated and hard-baked PR layer 460.

With reference to FIG. 6F, an additional PR layer 482 can be applied tothe silicon carrier wafer 462. In particular, the PR layer 482 can bespin-coated onto the carrier wafer 462 and patterned as a mask for deepreactive ion etching (DRIE). The PR layer 482 can comprise, for example,AZ9260. In certain embodiments, the PR layer 482 can be relativelythick. For example, the PR layer 482 can have a thickness of no lessthan about 30, 40, or 50 microns. In the illustrated embodiment, thethickness is about 40 microns. Etch selectivity between LCP and PR canbe approximately 1:1. Accordingly, the thickness of the PR layer 482, orPR mask, is desirably thicker than the combined thickness of the LCPsubstrate 410 and the carrier wafer 462 (which is slightly greater thanabout 25 or 30 microns, in the illustrated embodiment). Standardtechniques for high aspect ratio photolithography with thick PR can beused.

With reference to FIG. 6G, a hole, via, or opening 484 is created in thesubstrate 410 using an advanced oxide etching (AOE) DRIE system. Theopening 484 is provided to permit release of the sacrificial PR layer470, as discussed further below. The opening 484 is desirably alignedwith the opening 466 in the Ti/Au layer 464. The size of the opening 484can be selected as a trade-off between efficient release of thesacrificial PR layer 470 (larger openings 484 are desirable) anddegradation of the sensitivity of the substrate 410 to etching uponreduction of the sensing area (smaller openings 484 are more difficultto form). In various embodiments, the opening 484 can have a diameterthat is no greater than about 50, 60, 70, 80, 90 or 100 microns, or iswithin a range of from about 50 to about 100 microns or from about 80 toabout 100 microns. In the illustrated embodiment, the diameter is about90 microns. In the illustrated embodiment, the etching is performed forabout 90 minutes using low power (200W) at a low etching rate ofapproximately 0.3 microns/minute. Higher power would reduce the etchingtime, but the resulting high temperatures could cause the PR mask toreflow. As can be seen by comparing FIG. 6F to 6G, the thickness of thePR mask 482 has been greatly reduced by the etching process.

With reference to FIG. 6H, the assembly is submerged into PR stripper(e.g., PRS 2000) for a suitable period for removal of all of the PRmaterial that remains in the assembly—particularly the sacrificial PRlayer 470. In the illustrated embodiment, the assembly is submerged inPRS 2000 PR stripper for 48 hrs at 80 degrees Celsius to achieve the PRremoval. The assembly is then dried by a critical point drying (CPD)process.

With reference to FIG. 6I, fabrication of the pressure sensor 414 iscompleted by backside sealing of the openings 466, 484 in any suitablemanner. In the illustrated embodiment, Kapton® tape 488, which isavailable from DuPont Electronics, is used. The tape can be about 50microns thick. Thus sealed, the pressure sensor 414 includes adielectric region or air gap 448 that separates the capacitive plates444, 446. Again, it shall be understood that the dielectric region 448may be filled with any material (e.g., gas, liquid, solid, gel) whichallows the upper or lower plates 444, 446 to move or flex relative toone another, thereby creating a desired change in capacitance of thesensor 414.

FIGS. 7 and 8 show a top plan view and a cross-sectional view,respectively, of an illustrative embodiment of the pressure sensor 414that was fabricated via the method depicted in FIGS. 6A-6I. The sensor414 includes a diaphragm having a radius of about 250 microns, and fitswithin a volume of 500 microns×500 microns×100 microns. The pressuresensor 414 is suitably sized to be integrated on an implantable LCP tabfor placement in the anterior chamber of a mouse eye.

Testing of the pressure sensor 414 can be achieved by microprobingwithin a pressure-controlled chamber, such as within the probe station500 supplied by MMR Technologies of Mountain View, Calif., shown in FIG.9. The pressure range for the measurement of IOP of mice will normallyvary from 5 to 40 mmHg above atmospheric pressure. Accordingly, to testoperation of the pressure sensor 414 over a full range of standardoperation, the response of the sensor 414 in the range of 0 to 50 mmHgabove atmospheric pressure can be measured. The capacitance can bemeasured as the pressure is increased from low to high pressure using aninductance-capacitance-resistance (LCR) meter, such as a model 4284Ameter supplied by Agilent Technologies of Santa Clara, Calif. In theplot 510 shown in FIG. 10, the measured values were fluctuated in therange of ±1 femtofarads (fF) and the average value at a given pressurewas used. In particular, the plot 510 shows the average values of themeasured data as the pressure was swept ten times. The base capacitanceat atmospheric pressure was 1.167 f F. The capacitance increases nearlylinearly with an average sensitivity of 0.75 fF/mmHg. Accordingly,coupling the pressure sensor 414 with a suitable ASIC can permitmonitoring of IOP changes of less than 1 mmHg.

As shown in FIG. 11, the actual measured sensitivity of the pressuresensor 414 is in reasonable agreement with the theoretical or analyticalevaluations of the design—in particular, considering a residual stressof 35 MPa of the membrane, as calculated via Equation 1 above. FIG. 11shows a plot 520 that compares the actual data used in the plot 510 withtwo theoretical models of the pressure sensor 414: one of which assumesthat the membrane 442 has a residual stress of 0 MPa, whereas the otherassumes that the membrane 442 has a residual stress of 35 MPa. Theactual data closely tracks the theoretical model in which a residualstress of 35 MPa is assumed.

FIGS. 12A-12F depict another illustrative fabrication process that canbe used to create a capacitive pressure sensor 614 (FIG. 12). Specificdimensions, materials, and other features are provided with respect tothis illustrative process. While such specific dimensions, materials,and features may constitute independently patentable subject matter, itshould be understood that other suitable dimensions, materials, andfeatures may be used in fabricating other capacitive pressure sensors inaccordance with other embodiments of the present disclosure, and thussuch dimensions, materials, and features are not necessarily meant to belimiting.

The process in FIGS. 12A-12F utilizes direct parylene-to-parylenebonding. This approach can reduce the total thickness of the pressuresensor 614 and simplify fabrication, as compared with the pressuresensor 414 and the process of FIGS. 6A-61. For example, the process ofFIGS. 12A-12F eliminates the use of a multilayer structure that includessacrificial PR which is released to form the dielectric region or gapbetween the capacitor plates. It also eliminates backside sealing of thesubstrate (e.g., applying Kapton to an LCP substrate), thereby reducingthe overall thickness of the sensor 614. In certain embodiments, thedielectric gap can be encapsulated at a low temperature (e.g., about 230degrees Celsius) by a membrane transfer technique usingparylene-to-parylene bonding. This bonding process can be suitable forLCP substrates having a low melting temperature (e.g., about 315 degreesCelsius).

With reference to FIG. 12A, a piece of LCP 610 is attached to a thicksheet of copper 662, which can serve as a temporary carrier, via a layerof photoresist 660. The LCP can be of any suitable thickness, such asthose discussed above. In the illustrated embodiment, the LCP defines athickness of about 25 microns. The copper sheet 662 can have anysuitable thickness (e.g., about 635 microns in the illustratedembodiment), and may be obtained, for example, from K&S Engineering ofChicago, Ill. A mask layer 663 for reactive ion etching (RIE) can bedeposited on the LCP layer 610. For example, the mask layer 663 cancomprise a thin layer of Ti/Au, which may be deposited by sputtering.Any suitable thickness for the layer 610 is possible. In the illustratedembodiment, the layer 663 is about 60 nanometers thick.

With reference to FIG. 12B, the mask layer 663 is patterned by aconventional lithography process, and a cavity 667 is thereafter createdby an RIE process. A depth of the cavity 667 can be selected toultimately achieve a desired size of a dielectric region or gap 648(FIG. 12F). In the illustrated embodiment, the depth of the cavity 667is about 2 microns.

With reference to FIG. 12C, the mask layer 663 is etched out, and alayer 664 of Ti/Au is then deposited onto the LCP layer 610 andpatterned. The layer 664 can define one or more of a lower electrode 646and an electrical lead 622 (FIG. 12F). Any suitable thickness for thelayer 664 is contemplated. In the illustrated embodiment, the thicknessis about 0.5 microns. A thin layer of parylene 671 can then be coated asa bonding layer over the layer 664. In the illustrated embodiment, theparylene layer 671 defines a thickness of about 0.5 microns.

With reference to FIG. 12D, a sheet of copper 690 is coated with adetergent 692 (e.g., Micro-90®). Then, a parylene layer 672 is depositedon the metal sheet 690. The parylene layer 672 can have any suitablethickness, which may be selected based on the factors discussed above.In the illustrated embodiment, the thickness can be about 1.3 microns.The detergent 692 allows for easy detachment of the carrier sheet ofcopper 690 from the parylene layer 672 after the parylene layer 672 hasbeen transferred in a bonding process, as discussed further below.

With reference to FIG. 12E, the two parylene layers 671, 672 are bondedto each other using a compression molding press (e.g., any suitablepress available from Wabash MPI of Wabash, Ind.). Again, it shall beunderstood that the dielectric region 648 may be filled with anymaterial (e.g., gas, liquid, solid, gel) which allows the upper or lowerplates 444, 446 to move or flex relative to one another, therebycreating a change in capacitance of the sensor 414. In certainembodiments, the bonding may proceed at 230 degrees Celsius for 30minutes under atmospheric pressure with an applied force of 0.3 ton(U.S.). After bonding, the sample is dipped into a copper etchant, suchas ferric chloride. Due to undercut in the wet etching process, thecopper sheet 690 is easily separated from the transferred parylene layer672, even without etching the whole metal sheet 690.

With reference to FIG. 12F, another layer of Ti/Au 674 can be depositedon the parylene layer 672. The Ti/Au layer 674 can define one or more ofan upper electrode 644 and an electrical lead 620. The Ti/Au layer 674can define any suitable thickness (e.g., about 0.3 microns). Anadditional layer of parylene 676 can then be deposited over the Ti/Aulayer 674 and the parylene layer 672. In the illustrated embodiment, theparylene layer 672 is about 1.3 microns thick. A composite membrane 642portion of the pressure sensor 614 thus can include the Ti/Au layer 674sandwiched between the parylene layers 672, 676. As previouslydiscussed, the thicknesses of the various layers 672, 674, 676 can beselected to achieve the desired properties of the membrane 642, such asthe residual stress thereof. The LCP layer 610 can then be detached fromthe copper sheet 662 by removing the PR layer 660 via PR stripper. Forexample, in some embodiments, at least a portion of the assembly issubmerged in PR stripper for about 5 hours.

With reference to FIGS. 13 and 14, an illustrative embodiment of apressure sensor 614 was formed using the process discussed with respectto FIGS. 12A-12F. The sensing area is about 300 microns×about 300microns, and the thickness is about 30 microns. This thickness is anorder of magnitude lower than conventional silicon-based capacitivepressure sensors. FIG. 13 shows a 3-D image of the pressure sensor 614,and FIG. 14 shows a surface profile measured by a confocal laserscanning microscope (LEXT from Olympus). The membrane is almost flat andis positioned over a circular cavity of about 260 microns in diameter. Asurface roughness of the membrane is about 0.2 to about 0.3 microns.

Using the same measurement setup discussed above with respect to FIGS.9-11, the change of capacitance of the pressure sensor 614 was measuredin the pressure range of 0 to 50 mmHg above atmospheric pressure usingan LCR meter (Agilent 4284A) with the sensor 614 positioned inside thepressure chamber (MMR probe station). The pressure was swept ten timesand the average value was recorded. FIG. 15 shows the measured response710 of the sensor 614 to the change of pressure. The base capacitance is784.92 fF, and it increases with a nearly linear average sensitivity ofabout 0.3 fF/mmHg.

As shown in FIG. 16 (plot 720), the increase in sensitivity correspondswith a theoretical design that assumes a residual stress of 35 MPa forthe membrane 642. The sensitivity per unit area (milliimeter²) is about3.3 fF/mmHg, which is comparable to that of large-scale commercialsensor diaphragms, despite the much smaller area of the sensor 614. Thesensor 614 can thus be well-suited for IOP sensing with a resolutionless than 1 mmHg, such as when the sensor is combined with a suitableASIC.

In various other embodiments, the capacitance of the sensor 614 canincrease with a nearly linear sensitivity of no less than about 0.1,0.2, 0.25, 0.3, 0.4, 0.5, 0.6, 0.7, 0.75, 0.8, 0.9, or 1.0 fF/mmHg. Inother or further embodiments, the sensitivity per unit area(millimeter²) is no less than about 1.0, 2.0, 2.5, 3.0, 3.3, 3.5, 4.0,or 5.0 fF/mmHg.

As previously mentioned, in certain embodiments of pressure-sensingsystems (e.g., the systems 100, 300), the LCP substrate 610 on which anIC, an antenna, and a capacitive pressure sensor are integrated orotherwise coupled may bend so as to be accommodated within the limitedspace of the anterior chamber in the mouse eye 200. Based on thedimensions of a mouse eye 200 represented in FIG. 3, the radius ofcurvature on the posterior surface 212 of the anterior chamber 210 canbe calculated as about 4 millimeters. Such a small radius of curvaturecan be more challenging for pressure sensors, as compared with humaneyes or even other animal models. For example, the radius of curvaturecan be smaller for smaller eyes. With a smaller radius of curvature, thebending of the substrate can be more pronounced, such that it might havea greater effect on operation of the pressure sensor. In particular, thelower electrode 646 of the pressure sensor 614 may sufficiently thinand/or malleable so as to bend in conformity with the substrate 610.Such bending could interfere with pressure measurements; for example, ifthe sensing area of pressure sensor 614 is large, while the separationdistance between the capacitor's electrodes is small, the curvature ofthe lower electrode 646 could bring it into close proximity or evencontact with the upper electrode 644, in some instances. Curvature ofthe substrate 610 can also induce stress (residual stress) therein. Itcould also yield greater residual stress in the membrane 642. It can beestimated that a pressure-sensing system, which may also be referred toas an LCP tag, that is implanted inside the anterior chamber 210 of amouse eye has a radius of curvature greater than 4 millimeters, asindicated in FIG. 17.

With reference to FIGS. 18 and 19, in order to mimic the curved surfaceof an LCP tag that is bent when implanted, for purposes of testing, anepoxy mold 800 having a radius of curvature of about 5 millimeters iscreated. A pressure sensor 614 may be attached to the mold 800. Thepressure sensor 614 can be integrated into a substrate 610 having anextended area to facilitate testing of the pressure sensor 614.Moreover, the pressure sensor 614 can be electrically coupled withenlarged and extended contact pads 810 to facilitate the testing. Theassembled test sample is measured in a manner such as described above,such as by using the MMR probe station and the Agilent 4284A LCR meter.

FIG. 20 depicts a plot 830 that includes data obtained from a pressuresensor 614 that is operated with its substrate in a flat configuration,as compared with data obtained when the substrate is curved due to thepressure sensor 614 being attached to a curved-surface setup such asthat in FIGS. 19 and 20. The base capacitance and sensitivity on thecurved surface correspond with those on the flat surface. This resultshows that the estimated curvature of the implanted LCP substrate doesnot significantly affect performance of the capacitive pressure sensorhaving a sensing area of about 300×about 300 microns². Thepressure-sensing system is flexible, but the small size sensor is notaffected by this curvature and the resulting induced stress. A largerarea sensor, while being more sensitive, would have difficulty with thebending radius as the two capacitive plates could touch.

FIGS. 21A and 21B depict a further embodiment of a pressure-sensingimplant, or pressure-sensing system 900 that is configured to beimplanted in the eye of an animal. The pressure-sensing system 900 issimilar in function to system 100 described above and suitable forimplantation in the eye of a mouse, but with the addition of a lightemitting diode (LED) to provide a visual indication that power is beingtransferred, or the level of power transfer, between the externalradiation source and the system 100. In the illustrated embodiment, thepressure-sensing system 900 includes an integrated circuit (IC) 912, aMEMS pressure sensor 914, and an antenna 916, and an LED 917 that aremounted to or integrally formed with a substrate or tag 910. In certainembodiments, the LED output light intensity increases as the power beingreceived by the system 100 increases, thereby providing visualverification that the system 100 is operational. In other embodiments,the LED can be used as a communication medium, for example, by flashingbetween on and off states to convey a digital code or pattern in whichdata is encoded.

The integrated circuit 912 can be electrically coupled with each of thepressure sensor 914 and the antenna 916 so as to be able to electricallycommunicate therewith. In particular, separate portions of the pressuresensor 914 can be coupled with the integrated circuit 912 via a pair ofelectrical leads 920, 922, as discussed further below. The pressuresensor 914 can be configured to provide data (e.g., capacitance values)to the integrated circuit 912, which can be configured to store thedata, derive information from the data (e.g., derive calculated orestimated pressure values from the capacitance values), and/or deliverthe data and/or the derived information to the antenna 916 for deliveryto a remote data storage and/or analysis system. The integrated circuit912 may interface with the antenna 916, which in one embodiment can bean inductive loop antenna, in a similar fashion is described above withrespect to integrated circuit 112 and antenna 116 of system 100. In theembodiment of FIG. 21A-B, the antenna 916 may be a self-expandableNitinol loop antenna. The Nitinol provides a balance between antennaradiation efficiency and surgical feasibility. The pressure sensor 914and metal traces for interconnection between components are embedded inan ultra-thin form factor. Also in this embodiment, a double-sidedcopper (Cu) cladding 25 μm LCP (ULTRARAM 3850, available from RogersCorporation) is utilized as a carrier substrate, which makes the viaetching process faster compared to a thick silicon carrier substrate.

FIGS. 22A-F depict another embodiment of a capacitive illustrativefabrication process that can be used to create the capacitive pressuresensor 914. As shown in FIG. 22A, the back side Cu layer 905 of an LCPcarrier 962 is patterned for the etch mask of the reactive ion etching(RIE) process, and then the openings 984 through the carrier 962 arecreated by oxygen plasma etching during RIE.

With reference to FIG. 22B, the top side Cu layer 907 of the carrier 962is coated with 10 μm thick photoresist 909 to smooth the roughness ofthe Cu layer 907 on the carrier 962 and to allow for later release ofthe fabricated tag from the carrier LCP substrate 962. Then, a parylenesubstrate 911 of 10 μm is deposited on the photoresist 909 and a Ti/Au(2 μm ) layer 964 is sputtered and patterned, thereby creating metalpads 965 for mounting the diode 917 and antenna 916, and lower sensorelectrode 946.

With reference to FIG. 22C, to create the cavity 948 of the pressuresensor 914, a 4 μm sacrificial photoresist layer 970 is spun on thelower electrode 946 and covered with a 1 μm thick parylene layer 972.The metal pads 965 are then exposed by etching the parylene layer 972.

With reference to FIG. 22D, a thin Ti/Au layer of 0.1 μm thickness isdeposited and patterned to create the top electrode 944 of the pressuresensor 914.

With reference to FIG. 22E, before the Cu etching process, photoresist967 is coated to protect the top metal electrode 944 from the Cuetchant.

With reference to FIG. 22F, the Cu near the openings 984 and on thebottom of carrier 962 is then etched out, and a RIE process is performedto create the holes for releasing the photoresist 909, 967, 970 and etchout the outline of the tag 910.

With reference to FIG. 22G, the tag 910 is submerged into acetone torelease the photoresist 970 and create the cavity 948 of the pressuresensor 914. the cavity 948 is sealed with Kapton tape of 25 μm thicknessas shown.

With respect to FIG. 22H, a parylene layer 976 of 3 μm thickness iscoated for passivation so that the tag is compatible with the mouse eyeenvironment. The final diameter of the antenna is 2.2 mm and the area ofthe tag is 1.65×0.8 mm². The thickness of the thickest part of the tagis less than 100 μm. The functional tag and its implantation inside themouse eye are shown in FIG. 23.

In certain embodiments, to accurately determine the IOP, severalconversions occur to receive the data externally. These includeconversion from pressure to an electrical equivalent, digitization ofthe electrical equivalent, and finally transmission of that digitalequivalent. In one embodiment discussed herein, a pressure tocapacitance conversion is used to monitor pressure and to determine theaccuracy of the IOP monitoring system.

According to one embodiment, digitization of the capacitive datagenerated by the pressure sensor 914 is completed through frequencymodulation (FM). The measurement system uses a clock source that isdependent on the capacitor (initially a varactor) and second thepressure sensor. Initially using a varactor, FIG. 71 depicts how thefrequency output changes dramatically over large pF changes. The outputis nonlinear and the base capacitance of the sensor determines theoutput frequency as well as the resolution necessary to determine changein capacitance. The lower the base capacitance the larger the frequencychanges for given capacitance change. From data collected, a logarithmiccurve was fit to the data to determine the frequency based on capacitiveinput.

Freq(kHz)=−186.31n(Cap(pF))+726.46, R ²=0.981   Equation 8.19

This allows the conversion from capacitance to FM to have a higheraccuracy than a base capacitor using 10 pF.

Using a ring oscillator design, the FM modulated output signal rides onthe transmitted wave where it can then be demodulated and interpreted.Using the description in the materials and methods, a FFT of the outputsignal was taken from the pickup antenna. FIG. 72 depicts the sameoutput data obtained in FIG. 71. The same nonlinear logarithmic curve isobtained for the transmitted data, as what was created by themeasurement circuit.

From data described in Table 8.5 it is shown that there is no percentdifference of greater than 1% for the output data. Therefore there is anaccurate transmission of capacitive measurements transferred from themeasurement circuit to an external antenna, and that there is little orno loss of information when transferred.

TABLE 8.5 Comparison of FM output for measurement circuit andtransmitter Capacitance Measurement Ouput Tx Output (pF) (kHz) (kHz) %Difference 0.5 1380. 1369. 0.7859% ± 0.65% 1 1178. 1184. −0.4798% ±1.24%  5 523.0 527.8 −0.9242% ± 0.93%  10 324.7 323.9 0.2612% ± 2.15% 15230.8 229.4 0.6198% ± 1.59% 22 171.4 170.5 0.4967% ± 1.80%

Knowing that the base capacitance of the pressure sensor 914 isapproximately 1 pF, data suggests outputs around 1 MHz during thechanging of pressure. FIG. 73 depicts this data, and shows that themeasurement system is functional in determining frequency based on aninput pressure. Even though the capacitive change of the sensor is small˜0.5 fF per mmHg, the high frequency output of the measurement systemallows for a correlation of the Table 8.5.

As the pressure increases from this system, it is noticed that thefrequency changes negatively at a rate of 426 Hz per mmHg. Variousembodiments of the present disclosure having this resolution provideadequate correlation between frequency. The limit is put on the externalbase station, where there are no space and design constraints. To obtainthe resolution of 0.5 mmHg, the basestation should have a resolution of213 Hz. As the sensor increases in sensitivity, the output will have agreater frequency change, making it easier on the design of the externalbasestation.

The device used (FIG. 47), was developed for the need to monitor IOP ingenetically modified mice. This will assist researchers in determiningthe effects genes have on glaucoma and IOP changes. The device includesa 25 μm parylene substrate of dimensions 1.7mm by 0.8 mm. In oneembodiment, in this substrate, gold traces are created that connect anRF powering ASIC, a tuning shunt capacitor, and a red LED. The antenna,a 2.5 mm 2.5 mm nitinol loop may be gold coated and attached at the headof the substrate. In one embodiment, the RF powering ASIC is similar tothose disclosed by Chow, et. al. in “A Miniature Implantable RF-WirelessActive Glaucoma Intraocular Pressure Monitor,” Institute of Electronicsand Electrical Engineers Transactions on Biomedical Circuits andSystems, Vol. 4, No. 6, December 2010, pp. 340-349, the entire contentsof which is hereby incorporated by reference in its entirety.

After the initial anesthesia has taken effect, the animal was mounted ina stereotaxic frame. A mouse-specific nose cone was used to maintain ananesthetic plane within the stereotaxic frame. Throughout the procedure,sterile saline was applied to the operated eye at regular intervals toavoid drying. After checking for adequate anesthesia (e.g., toe pinchand observation of vitals via the pulse oximeter), a trochanter was usedto puncture the eye at about 3-5 mm from the corneal limbus and implantthe intraocular pressure monitoring device into the anterior chamber.The device is roughly 300 microns cubed. The trochanter was removed andthe insertion site was sealed using a thin layer of adhesive (DermabondTopical Adhesive). Once the adhesive has dried, recordings were taken.

A secondary insertion technique is also viable. The eye was washed withsterile PBS. A stab incision of no more than 1 mm was made at about 3-5mm from the corneal limbus using a 3 mm microsurgical blade, or similardevice. A volume of Viscoat Intraocular Viscoelastic Injection (or othercommercially available alternative) was used to form the anteriorchamber to allow for better manipulation of the sensor, while minimizingtrauma to the eye. (Viscoat Intraocular Viscoelastic Injection, orsimilar solution, is sterile, non-pyrogenic, transparent viscoelasticpreparations of a highly purified, noninflammatory, high molecularweight sodium hyaluronate or similar substances. These substances areroutinely used in anterior segment ophthalmic surgical procedures inhumans. It coats the iris, posterior corneal surface and anterior lenscapsule, which helps protect these tissues from injury during thesurgical procedure.) The intraocular pressure monitoring sensor wasinserted through the incision, the intraocular viscoelastic injectionsyringe removed, and the incision sutured closed. Finally, a topicalantibiotic, Vetropolycin, was applied to decrease inflammation for theremainder of the surgery.

The powering measurement setup (FIG. 50) was set to remove many factorsincluding device location, monitoring offset, etc., and have only twofunctional factors for testing; these are frequency and power. In orderto power the LED device an Agilent N5182A MXG Vector Signal Generatorwas used to create the initial signal. The signal was passed through anOPHIR 5161 RF power amplifier, and sent to a Dorado AN-GH1-12S hornantenna. To collect data, video recordings were taken in a black room torecord the intensity of red light emitted from the diode. A camera wasset at a specific distance from the LED device.

Powering levels from −15 to 3 dbm (25-43 dbm at horn antenna) outputfrom the signal generator and frequency ranges from 1-2 Ghz were tested.A Labview code was developed in order to randomly choose a frame formthe video that corresponds to each frequency and power. These pictureswere then imported into Matlab for further analysis.

Given a specific input voltage, the LEDs have a corresponding currentpower relationship given by the manufacturer. This gives a conversion ofLED intensity area to power input. Second, the device is tested ex-vivo.These devices were tested using the described power and frequencylevels, but were tested in air at specific distances from the hornantenna and camera. Finally, the devices were implanted into mice eyes.This determined light intensity created by the same powering schemes asdescribed ex-vivo. The animals were tested weekly after implantation tounderstand the effects the animal, inflammation, and time have on thepower efficiency received by the LED device.

Mouse eyes were collected by first excising some of the surrounding skinand muscle using fine scissors and hemostats. A small, blunt spatula wasthen carefully inserted lateral to the eye against the zygomatic bone tothe back of the eye. Moving the spatula against the bone, at leastpartly around the eye and eventually to the back wall of the orbital,was then performed to sever muscle and nerve connections. Eyes were thenstored in HBHS in Eppendorf tubes at 4° C.

For histological imaging, eyes were secured in 1% agarose in PBS on acoverglass-bottomed dish. Laser scanning confocal microscopy wasperformed using on a Zeiss LSM10 using a translating microscope stage.Dil or DiD fluorescences in vasculature were imaged using the 543 nm or633 nm laser lines, respectively. Reflectance of 633 nm light from theimplanted device and the tissue surface was also collected, to capturethe device location relative to vasculature.

The size of the mouse anterior chamber has less than 300 um of workablethickness while the total diameter of the eye is approximately 2.5 mm.Therefore a device with a maximum diameter of about 2 mm and a thicknessof about 300 μm is used in some embodiments of the present invention.Initial work using the FR-4 and LCP designs worked towardsminiaturization of an implantable device (FIGS. 51-54). FIGS. 51-54 showthe form factor change and miniaturization to get a LCP LED devicecreated within size constraints.

Various embodiments of the present disclosure perform well in ex-vivo,and yet other embodiments of the present invention pertain to use in theoptic electrode design for neural stimulation. Initially, FIGS. 56-58,it was conceived that the devices could be created on LCP and sealedhowever, due to constraints of less than a 800 μm incision for the eye,adaptations were made to the overall design. various embodiments of thepresent invention use nitinol or other biocompatible shape-memorymaterials as the antenna for the mouse IOP device. Using the memoryalloy properties of nitinol various embodiments contemplate apre-implanted shape that is able to squeeze the implant through aspecially made inserter, or by hand, through the incision and allow itto come to full size once inside the eye (FIG. 58). Using this initialdesign concept of a nitinol loop antenna, an LED device was created. Itis understood that the LED device is utilized to prove the functioningof the device after it has been implanted (FIG. 55).

In order to properly test the system in a mouse, full system devicesneeded to be created that fit inside the limits of mouse eye anatomy.One embodiment of the present invention includes a 300 μm full systemdevice to accomplish this goal. This ASIC is composed of four pads, twofor the RF powering and two for the MEMS capacitor. This full systemwill then connect to the substrate and tested for full systemfunctionality (FIG. 56). This system again does not have the sensorattached, but power capabilities are demonstrated

Analysis of the LED on the implantable device (e.g., system 900) ishelpful in determining the power transfer between the 2.5 mm loopantenna and an external source. Using the current voltage relationshipdictated by the LED manufacturer allows for a power conversion based oninput voltage. This input voltage then relates to the area displayed bythe LEDs intensity.

Table 8.3.1, shows how current and power consumed by the LED devicerelates to input voltage for a particular LED supplied by CREE, Inc. ofDurham, N.C.

TABLE 8.3.1 CREE LED voltage current relationship Voltage (V) Current(mA) Power(mW) 1.625 1 1.625 1.7 3 5.1 1.8 8.5 15.3 1.9 17.5 33.25

Plotting this data, a curve fit equation can be used to relate voltageand current. This leads to an understanding of power consumed by the LEDduring powering using the relationship for power. For a given inputvoltage the LED lights up to a specific intensity. The LED intensity cantherefore be monitored to determine the power being received by thesystem. In a dark room, a camera positioned a specific distance awayfrom the LED device captures the LEDs intensity. By collecting thisdata, LED output intensity can be related to power consumed by the LED.Once the power being consumed by the LED is known, the power beingconsumed by other components within the system 100 can be determinedbased on the known electrical relationships between the other componentsand the LED. Thereafter, the relationship can be used to determine powerbeing received and/or consumed by the system after implantation.

Each device, although designed for the same specifications, has adifferent resonant frequency that makes it efficient. This is useful inRF powering. In certain embodiments, there is little to no room formatching networks on the implanted chip. In one example embodiment of apressure sensor system similar to system 900, the device tested wasfunctional in the 1-2 GHz range (FIGS. 60 and 61) to determine thefrequencies which provide the most efficient power transfer. Anothertested device was found to be functional at around 3.0 GHz (FIGS. 62 and63).

FIGS. 61 and 63 show a power consumption of 2 mA and 1.5 mA for thefirst and second tested devices respectively. Using this data arelationship of power input (power at the antenna) and the powerconsumed by the device is observed. In one example test, withapproximately 8 watts of power being transmitted to the implanted devicefrom the external receiver, at the peak frequencies the LED is consumingan amount of power which correlates to a consumption of 1.5 mW.

The devices tested have shown proper functionality after implantation.In terms of proper functionality and power consumption, one device wasable to light up at the same distance from the external powertransmitter as the ex-vivo testing. FIG. 59 depicts a mouse with theimplanted LED device. In the visible eye the implant can be seen. FIGS.64 and 65 show a relation of the power consumed by the LED and powerinput at a date 1 month after implantation.

During testing of the above devices, differences in the in-vivo andex-vivo results were observed. First, the power consumed given the sameinput power decreased, and the frequency at which the LED is poweredjumped. Upon analysis of power consumed, at its peak, the ex-vivo devicedisplayed a power consumed 2.00 mA ex-vivo at 1.6 GHz while in-vivo hada 1.01 mA power consumed at 1.8 GHz.

This power decrease is expected under the open condition (air), andimplanted condition (anterior chamber of mouse eye). One reason is thatthe boundary conditions set up by the anatomy of the mouse; cornea,aqueous humor, and skeletal structure causes reflections of the energyat each boundary. Finally, as the remaining energy reaches the implanteddevice, a decrease in power consumed is observed for the LED device, andtherefore a 50% difference in power consumption between ex-vivo andin-vivo analysis.

The second observation is with the resonant frequency where the LEDconsumes the most energy. For the first animal (FIGS. 60 and 61) andimplanted data (FIGS. 64 and 65) there is a jump of 200 MHz. This isobserved because the antenna for RF coupling is not directly matched toany given frequency. Its original harmonic includes the antenna itselfsetting its resonant frequency at approximately 1.6 GHz. When the deviceis implanted inside the animal, the medium around that antenna willchange that resonant frequency based on its electrical characteristics.

Two other animals beside the one described also were able to pick uprough LED data, not using the LED intensity capture method, but byplacing the mouse inside a horn antenna used to receive the signaloutput by the implanted device. Gathering these measurements consistedof changing the frequency and power while recording the frequency whichpermitted the largest LED intensity by visual inspection. Table 8.3.2depicts the data received. With regards to device 3, following the samescheme as the previous 2 devices in Table 8.3.2, an increase infrequency is expected. However, this was not the case. Devices 1 and 2were of the same initial batch of fabricated antennas, therefore theyhad a similar RF powering scheme. The third device was of a differentrun, and due to the connections had a different characteristic resonantfrequency, at approximately 3.1 GHz. Following implantation of theanimal, data collected at the first and second weeks has shown thefrequency has moved, not up as in the first set of devices, but down tothe approximate 1.8 GHz data point observed in the first and secondmice.

TABLE 8.3.2 Resonant frequency to obtain brightest LED intensity Mouse 1Mouse 2 Mouse 3 Ex-Vivo Functioning  1.6 GHz 1.38 GHz  3.1 GHz FreqIn-Vivo Functioning ~1.8 GHz ~1.8 GHz ~1.8 GHz Freq

This shows that the medium in which a passive device is implanted has aneffect on its output LED response. This is helpful in showing thatdifferent animal models and different anatomical locations exertdifferent natural frequencies (i.e. anterior chamber of eye, surface ofskin above skull, and inside artery). Following implantation times ofone, two, and four weeks, histology was conducted on the eyes of theimplanted animal, discussed below in Example 3.

Referring to FIGS. 36 and 43, various embodiments of the presentinvention comprise a capsular tension ring (CTR) form 990. As shown, thecapsular tension ring 990 has a base portion 992 for housing the sensorand related electronics and two curved arms 994 extending from the baseportion 992. This device will already have curvature built into themechanical integrity of the device, and allows the device to maintainthe position of the sensor within the eye. FIG. 43 shows such a deviceafter implantation in an animal eye. FIG. 35 shows a further embodiment,where the base portion 992 is offset, leaving a single curved arm 994extending from the base portion. In certain embodiments, nitinol makesup the dipole antenna arms 994 of the CTR design.

Various embodiments of the present invention include a base station 1000for powering of the pressure sensor circuitry, and also for receivingdata from the pressure sensor (referring to FIG. 37). This base stationshould have universal functionality so that it has the ability to workwith any testing facility and researcher or doctor. The two initial basestation designs were for functionality of glaucoma for both the humanIOP and mouse devices. The mouse base station was made to fit under theshelving where mice would be held so that it would not interfere witheveryday activities. For the human device it was determined that theophthalmologists want the antenna part of the device close to thepatient, but the equipment to run the system farther away from thepatient and near the computer and equipment that they use.

The systems described above have gone through functionality testing.From the work described previously along with silicon 6 mm by 3 mmdevices compared implantable powering of the LTCC loop and SI loop atspecific distances and differing RF powers. Upon implantation into therabbit anterior chamber RF power received was compared between LTCC andSI devices (FIG. 46). What this shows is when the systems increased indistance from its radiating source that more power was needed toattribute the same current consumption by the chips. This demonstratesthat RF powering had functionality for wirelessly powering devices.Also, the method of capturing the LED intensity and/or power consumptionfor a given device can help improve the understanding of the energybeing coupled to a device while removing the wires that are normallyused to analyze energy transfer.

In certain embodiments, the circuitry within the pressure sensing system100, 900 is configured to provide for signal processing of thecapacitance signal. In one embodiment, the pressure sensor interfaceswith circuitry that provides the sensor signal at a higher orderharmonic, such as the 3^(rd) order harmonic, of the radiowave frequencythat was utilized to power the sensor. In still further embodiments, thecircuitry includes a low-power, 10 bit analog to digital converter thatprovides a high resolution signal of the changes in capacitance. Highisolation between the fundamental and harmonic tone and higher antennaefficiency at the harmonic frequency can therefore be achieved.Furthermore, due to the high isolation, detecting a low power harmonicsignal in the presence of a relatively high power transmitted toneresults in alleviation of the dynamic range constraint of the receiver.

Moreover, as the 3^(rd) order scheme enables three times more frequencyshift per unit capacitance change than using the fundamental tone,higher resolution measurement of a pressure change can be accomplished.Also, compared with the traditional inductive coupling method, harmonicdetection provides further sensing distance since the coil is utilizedas an antenna rather than an inductor.

In some embodiments, the inductive Nitinol loop antenna and the MEMScapacitive pressure sensor form an LC resonator circuit. When thecapacitance of the MEMS sensor changes because of IOP variation, thefundamental tone resonance frequency of the LC resonator circuit isshifted and results in a change of the resonance frequency of the 3rdorder harmonic signal being transmitted back to the external receiver.This shift is then measured to determine the change in pressure beingsensed by the MEMS sensor.

A system 2400 for measuring intraocular pressure according to oneembodiment is shown in FIG. 24. First, a signal generator 2402 (e.g., amodel N5182A by Agilent) which generates the fundamental tone signals isconnected to a transmitting patch antenna 2404 through low pass filter2406 and a power amplifier 2408. In the illustrated example, the poweramplifier 2408 has a gain of 42 dB (5161 RF power amplifier by OPHIRRF), however it shall be understood that higher or lower gain values maybe used depending on the particular application. In order to receive thereradiated 3rd order harmonic signal from an implanted pressure sensortag 2410, a horn antenna 2412 is connected to a spectrum analyzer 2414(e.g., a model E4408 by Agilent) via high pass filter 2416 and a lownoise amplifier 2418. The mouse 2418 with the 3rd order harmonic taginside its eye is anesthetized and placed on a heated surgical table. Inthe illustrated example, the transmitting patch antenna 2404 and thereceiving horn antenna 2412 are placed at distances of 5.5 cm and 11.5cm, respectively, away from the mouse eye. The distances are chosen tomimic the distance between the antenna and a mouse moving in a standardcage (30×30×15 cm³).

The frequency of the fundamental tone being output by the signalgenerator 2402 is swept from 2.2 GHz to 2.7 GHz, and the 3rd orderharmonic signal is monitored from 6.6 GHz to 8.1 GHz on the spectrumanalyzer 2414. In the illustrated example, the total power transferredto the patch antenna is 23 dBm. The pressure within the mouse eye isincreased in 10 mmHg increments from 20 mmHg to 40 mmHg aboveatmospheric pressure, with the pressure kept constant for at least 10minutes. IOP in mouse strains often varies between 5 mmHg to 40 mmHgunder normal and disease conditions, with an IOP above 20 mmHgconferring increased glaucoma risk. The measurement range of 20 mmHg-40mmHg is adequate to identify high IOP in some embodiments associatedwith glaucoma.

The collected data is fitted using a shape-preserving interpolantfunction in MATLAB. The resonance frequency of the 3rd order harmonicsignal appears between 6.68 GHz and 6.73 GHz with the change in IOPbetween 20 mmHg to 40 mmHg. FIG. 25 shows the shift of the resonancefrequency of the 3rd order harmonic signal. The data is normalized withrespect to the maximum power level of each sweep. The average frequencyshift of the 3rd order harmonic signal per unit pressure isapproximately 1.5 MHz/mmHg, which is ultimately capable of monitoring anIOP variation of 1 mmHg from a mouse in a cage using the samemeasurement setup.

In some embodiments of the present disclosure, the sensor assembly(e.g., system 100) includes means for resolving the change in IOP. Adifferential two-stage amplifier according to some embodiments operateswith a MEMS capacitive pressure sensor. It compares the value of thecapacitance of the MEMS sensor with an internally adjustable referencecapacitor and gives a differential output voltage to represent thedifference. A correlated double sampling function is implemented tosuppress the 1/f noise from the input transistors in the front-endamplifier, the offset voltage due to the amplifiers, and also errors dueto the switches in the amplifier.

The capacitance measurement in some embodiments is taken by toggling thevoltage on the common node between the two capacitors with a stepvoltage of value VS. This causes a differential voltage on the output ofthe pre-amplifier equal to VS(CM-CR)/CI. The value of CI is 180fF.

There can be an input common mode adjust circuit, which compensates forthe fact that the step voltage VS may alter the input common modevoltage. By adjusting the input common mode voltage in this way, theoffset due to parasitic input capacitance mismatch can be minimized.

The clock generator gives four clocking signal for the amplifiers. Theycan be seen in the corner of FIG. 74. From these clocks, three phasesare generated. The first is the reset phase where the capacitors arezeroed and the outputs are set to the common mode voltage level. In thesample_1 phase, the noise, including the 1 /f noise is sampled onto thecapacitors CH. Then in the sample_2 phase, the capacitor measurement istaken. Since the noise is relatively low in frequency compared to thesample rate, it has the same value between the sample_1 and sample_2times and is cancelled. The total gain at the output is (CH/CII)(VS(CM-CR)/CI). The nominal base capacitance of the sensor can be variable.This variability is compensated for by dynamically adjusting theinternal reference capacitor. When the supply voltage vdda is appliedand the power-on-reset fires, the initial value of the referencecapacitor will be relatively low compared to the MEMS capacitor. As eachsample is taken, this value will be incremented with the counter untilthe value of reference capacitor is close to but slightly higher thanthe reference capacitor. When this happens, the comparator will fire andthe value of the capsel bus coming from the counter will be latched. Theclock signal vcsample is used for timing the comparator samples.

Some embodiments of the present invention include an amplifier whichincludes some or all of the following: 1/f noise suppression using CDS;tunable reference capacitor to allow for variation in the MEM base capvalue; an input common mode adjust circuit to minimize output offset dueto parasitic input capacitor mismatch; and low power (average current<50 uA). The following table presents specifications for an amplifierfor amplifying a pressure signal according to some embodiments of thepresent invention:

Block Parameter MIN TYP MAX UNITS CDS C-V Power Consumptions (current)50 uA Sample Rate 2 kHz Pressure Sensitivity 0.5 mmH Adjustable Base Cap12.5 fF/bit No. Bits accuracy for pressure 9 . . . Bits reading

Various embodiments of the present disclosure pertain to the use of ananalog to digital converter having more than about ten bits ofresolution. A system according to one embodiment of the disclosure isshown in FIG. 26. Various aspects of the analog to digital conversionare shown in the following table:

Specification Targeted Achieved Note Process XFAB 180 nm Voltage 1.8 V+/− 1.8 V +/− Core devices for 10% 10% analog design Current <50 uA 30uA Preliminary estimate Mode Fully Fully Better CMRR differentialdifferential and 2x dynamic range Input Signal 5 mV to V 5 mV to range1.2 1.2 V Input Signal 1 KHz 1 KHz Bandwidth Internal 250 MHz 250 MHzStable clock Clock source Resolution 10 bits 10 bits ENOB = 9 bits INL<1.5 LSB <1 LSB DNL <1 LSB <1 LSB OP Code Signed 0111111111 = +1.2 V;Magnitude 1111111111 = −1.2 V Comparator <0.5 LSB <0.5 LSB OffsetComparator <0.5 LSB <0.5 LSB Designed to Resolution begreater than 10bits Phase margin >55 deg 72 deg of Preamp Power <100 uW 60 uW Area <350um × 320 um × 300 um 240 um

One eye pressure sensing system 2600 is shown in FIG. 26. The systememploys a variable MEMS pressure sensor 2602, similar to sensors 114,314, 414 to sense real-time eye pressure waveforms. The sensor changesits capacitance based on pressure applied to it and is processed by anon chip sensing circuit CDS-CV (Coherent Double sampling—Capacitance toVoltage converter) 2604, encoded by a 10 bit successive approximationregister analog to digital converter (SAR ADC) 2606 and wirelessly sendthe data out by a transmitter 2608. Wireless RF signals from an externalsource (not shown) are captured at block 2612 and fed to powermanagement block 2614 for powering the system 2600.

System 2600 preferably includes an ultra low power 10 bit SAR-ADC 2606along with stable clock source 2610 to encode eye pressure signalwaveform. FIG. 27 shows the top level schematic for an example 10 bitSAR ADC 2606. It includes the following blocks: (a) comparator 2702; (b)capacitor array (DAC and S/H) 2704; (c) delay elements and drivers 2706;(d) SAR control logic 2708; and (e) switches 2710. Discussion of theseblocks is provided with reference to FIGS. 28, 29, 30, 31, 32, 33, and34.

Each A/D Converter contains at least one comparator 2702. A comparatoritself can be considered a 1-bit A/D Converter. In the presented ADCdesign, the comparator plays a key role; it should be able todiscriminate voltages as small as 878 uV. FIG. 28 shows the top levelschematic for a comparator 2702 according to one embodiment.

A helpful specification is the offset voltage. The offset voltage shouldbe smaller than 0.5 VLSB=878 uV. Clearly, to reach this requirement,offset-cancellation techniques should be applied. In fact, differentialamplifiers, as the ones used in the pre-amplifiers 2802 of thecomparator have an offset voltage of 1mV to 10 mV if they are realizedin CMOS technology. In practice, the residual offset after performingdynamic offset cancellation can be in tens of micro-volts.

In the illustrated embodiment, 3 amplifiers are employed for outputoffset storage and auto zeroing with 22 dB of gain. Preamps designed tohave low offset. Each preamp is consuming 4 uA of current. The gain of asingle preamp cannot go beyond 25 dB in output offset storage (otherwiseamplifier output will saturate for high offset values), therefore 3preamplifiers 2802 are used, each one if having 22 dB of gain. Anexample schematic for single preamplifier is shown in FIG. 29.

An auto zeroing clock scheme is shown in FIG. 30. A non-overlappingclock signal is generated for auto-zeroing and offset cancellation.Sample and hold cycle and three auto-zero phase timings are a usefulconsideration in offset cancellation. The auto zero phase starts after afinite time delay (2 ns as one example) in hold phase of sample and holdcycle.

In order to establish full logic levels and synchronize the instant adecision is taken with other blocks, the back-end of the comparator 2702includes a latch 3102. The output nodes outp 3104 and outn 3106 (seeFIG. 31) are pre-charged to when the clock is low. To prevent staticcurrent flow through the two branches of the latch, a pMOS transistorcontrolled by clk_cuts off the cross-coupled inverter pair during thepre-charge phase. The amplified signal (output of preamplifier stage 3)is applied to the latch through the middle two pMOS transistors, whichprovide an additional gain. The pre-amplified signal can thereforeovercome the offset voltage of the latch.

One implementation of SAR ADCs uses a binary weighted capacitor arrays,however, area and power increase exponentially with resolution. In fact,N-bit resolution requires 2^(N) unit capacitors.

Split capacitor arrays as well as C-2C ladders reduce the totalcapacitance, reducing area and power. However, the parasiticbottom-plate capacitance of series capacitors affects the linearity ofthe ADC. If the ratio of bottom-plate capacitance over nominalcapacitance is precisely known, this non-linearity problem can be dealtwith during the design phase by scaling some of the unit capacitors.Another approach includes shielding the series capacitors.

XFAB18.0 CMOS technology comes with a Metal Capacitor (MIM) module, alsoreferred to as Metal Insulator Metal (MIM) capacitor module. For a lowerbottom-plate parasitic capacitance, the MIM module can be insertedbetween the second-last and last (top) metal layers.

The unit capacitor sizing should consider KT/C noise, 10-bit accuratematching, timing and power consumption. To decrease power consumptionand increase speed, the unit capacitor should be as small as possible.On the other hand, to improve MIM capacitor matching, noise immunity andconsequently the ADC's accuracy, the unit capacitor should be as big aspossible. One embodiment of the present disclosure includes a unit capvalue as 84 fF, to consider trade-offs. The cap array is shown in toplevel schematic in FIG. 27.

The Successive Approximation Register (SAR) sets the switches—as afunction of the current state of the conversion and the comparator'sresponse—and stores the digital output code to be issued at the end ofthe conversion. The main building blocks of the SAR control logic are:(a) 4 bit counter; (b) output register; and (c) combinational network.

FIG. 32 shows the SAR control logic and its sub blocks according to oneembodiment. The 4 bit-Counter counts from 0 to 16, thereby cyclingthrough the 10 states (some embodiments use 10 states out of 16available) of one conversion. The counter is controlled by the 250 KHzreference clock or clk. The count or state of the counter changes oneach positive transition of clock clk.

The combinational network calculates the control commands (Set, Resetand Select) for the output register bank, indirectly setting theswitches, accordingly to the history of the current conversion. Theoutput registers store the correct position of the switches for thecurrent conversion and contain the output code at the end of theconversion. They are thus multifunctional. Their multi-functionalitysaves registers. The output registers eventually change their content ona positive transition of clk1, which is delayed with respect to clk.

The switches in this design are analog transmission gates. Cap arrayswitch analog multiplexer and transmission gates are useful because oflow and constant Ron requirement. FIG. 27 shows the analog Switchesalong with cap array.

As one example, a test of an at least partly differential sinusoidalsignal of 400 mV peak to peak is applied to ADC input. Top level testbench is includes 10 bit SAR ADC, one shift register and one ideal DAC.The top level simulation results for ADC and Cap array DAC and sampleand hold circuit are shown in FIG. 33.

A top level layout (without pads) is shown in FIG. 34. The top left andtop right blocks are the cap array and SAR logic respectively. Thecomparator and S/H switches are shown on the bottom.

Pressure-sensing systems such as described above can be used in avariety of contexts other than IOP monitoring in mice or other animals.Numerous medical procedures and applications can benefit from extremelysmall pressure sensors that are biocompatible, flexible, and/ornoninflammatory. For example, some embodiments may be used within theheart of a subject (e.g., a human) in any suitable procedure orapplication. The pressure-sensing system may include an active device(e.g., an integrated circuit) that is coupled with a passive device(e.g., a pressure sensor) that passively measures fluid pressure in theheart. The active device can transfer data out of the body to equipmentconfigured to receive, store, and/or process the data.

Other or further embodiments may be used in the context of nephrology.For example, certain dialysis applications can benefit from sensor suchas described herein. In some embodiments, a pressure-sensing system canbe coupled with (e.g., integrated into) a catheter, such as a centralvenous dialysis catheter. The pressure-sensing system may be positionedat an outer surface of the catheter. The system can be used to measureblood pressure during dialysis and ensure that the pressure does notdrop below a certain safety level. The sensors thus can be used toprevent the triggering of cardiac arrest.

Other or further embodiments can be used to track the pressure ofcerebral spinal fluid. For example, pressure-sensing systems such asdisclosed herein can be implanted in the skull so as to monitor theonset or progression of hydrocephalus.

Still other or further embodiments can be used as monitoring devices forbreast implants. For example, a silicone breast implant could include apressure-sensing system. The pressure-sensing system may be freefloating within the breast implant or attached to an inner surface ofthe breast implant. The placement of a silicone breast implant in thebody (especially when it's submuscular, e.g., beneath the pectoralismajor) creates a certain pressure within the breast implant that exceedsthe pressure outside, since the surrounding body tissue has acompression effect on the implant. The pressure-sensing system can beconfigured to transmit data regarding the fluid pressure within theimplant to remote equipment. If a rupture occurs, the pressure withinthe breast implant will drop as the implant fluid (e.g., silicone orsaline) leaks out and the volume of the implant decreases. The drop inpressure is detected by the pressure sensor and/or remote equipment.Such a system can advantageously provide an early warning system forimplant rupture, and can reduce the number or frequency of checkupvisits for patients at which expensive MRIs are generally performed.

Yet other embodiments include the use of the disclosed pressure sensorsin a cranial cavity (such as for real-time monitoring of cranialpressure), and implantation in a spinal cavity (for real-time monitoringof the pressure of spinal fluid).

It is contemplated that the test methodologies used in the belowexamples and others disclosed herein may be used to test any embodiment.

Example 1 Biocompatibility Testing

Various embodiments of the present disclosure can include various typesof coating of the implantable device. Additional tests include aparylene coating often used to ensure biocompatibility for implantabledevices. The effect of coating a low temperature co-fired ceramic (LTCC)with parylene was explored and compared to quantify the possibleimprovement of biological tissue response. Testing was also done onalumina, which is used as a baseline material for biocompatibilitycomparisons.

Biocompatibility was verified through in-vivo trials performed on 6 NewZealand White Rabbits. This animal model was chosen in accordance withISO standards for biocompatibility testing. The comparative study donein this work focuses on ISO 10993-part 6, which delineates tests forlocal effects after implantation including inflammatory response andfibrous encapsulation.

Surgical procedure follows the Purdue Animal Care and Use Committee(PACUC) approved protocol (PACUC No. 08-004) beginning with a one-weekacclimation period prior to surgery. Surgery begins with the injectionof an anesthetic induction solution comprised of ketamine and xylazine.After anesthetization, verified by a toe-pinch test, anesthesia wasmaintained by an intravenous (IV) drip of Propofol. Prior to incision, alocal anesthesia, Zylocaine, was injected at each implantation site.Incisions were made along the dorsal side of the rabbit and deepenedinto the muscle tissue. The alcohol-sterilized materials were implantedinto the muscle and the incisions are sutured and stapled to seal thelayers of muscle, subcutaneous tissue, and dermis. The rabbits weregiven analgesics and triple antibiotic ointment as needed.

Histological examination was used to evaluate the inflammatory responseand fibrous encapsulation. After implantation durations of 1 week, 2weeks, and 4 weeks, the animals were euthanized through an injection ofsodium pentobarbital solution. The implant along with the surroundingtissue was excised. The extracted samples were embedded in paraffin waxand sliced into 50 μm thick sections using a vibratome. The tissue wasstained to provide contrast using a solution of Mayer hematoxylin pairedwith eosin. Specifically, eosin was used to highlight the elastic andreticular fibers while hematoxylin targets the nucleic acids andergastoplasm. Optical microscopic examination was then performed on thestain-enhanced tissue slices and the fibrous encapsulation isquantified.

Initial inspection evaluating inflammation showed minor redness withoutuniform density in all materials through the first 2 weeks of the invivo studies. After the 4-week implantation period, DuPont™ 951 andalumina still had visible minor redness while there was no visibleinflammation in the other materials used in the study.

To quantify the results, histological analyses of tissue slices wereperformed and fibrous encapsulations measured at three different pointsaround the implant site for all 6 rabbits. Optical micrographs of thematerial cases with the good tissue response, LCP and parylene coatedLTCC, are shown in FIG. 38. Thickness measurements at thetissue-material interface are averaged and the data is tabulated inTable 6.2.1. The different columns represent the measurements taken atvarious implantation durations of 7 days, 14 days, and 28 days.Post-surgery healing and corresponding inflammation influence the day 7and 14 measurements. After 28 days, the lack of post-surgery healingeffects allow for more precise measurements of the biological tissuereaction in the form of fibrous encapsulation.

TABLE 8.2.1 Material Encapsulation Thickness Encapsulation EncapsulationEncapsulation after 7 days after 14 days after 28 days Material (mm)(mm) (mm) DuPont ™ 951 0.223 0.026 * DuPont ™ 943 0.336 0.012 0.018Heraeus HL2000 0.070 0.019 0.059 Parylene Coat 0.201 0.015 0.000 LCP0.046 0.005 0.000 Silicon 0.098 0.004 0.000 ACA 0.185 0.059 0.017Alumina 0.081 0.020 0.036 * Data could not be extracted due to slicingerror

The data for all the tested materials are compared with that of thebaseline, alumina, and the differences are tabulated in Table 6.2.2.Each column represents the thickness differences in comparison with thealumina case for the corresponding implantation duration. A plus signrepresents more fibrous encapsulation than alumina while a minus signrepresents less encapsulation; the measurements plotted in FIGS. 39-41.FIG. 39 show a comparison between the materials after an implantationduration of 7 days, FIG. 40 shows 14 days, and FIG. 41 represents theresults after 28 days. The plots show the averages of the multiple datasets and the corresponding standard deviations.

TABLE 6.2.2 Encapsulation thickness with respect to control (alumina)Encapsulation Encapsulation Encapsulation after 7 days after 14 daysafter 28 days Material (mm) (mm) (mm) DuPont ™ 951 +0.142 +0.007 *DuPont ™ 943 +0.255 −0.008 −0.018 Heraeus HL2000 −0.011 0.000 +0.023Parylene Coat +0.120 −0.005 −0.036 LCP −0.035 −0.015 −0.036 Silicon+0.017 −0.015 −0.036 ACA +0.104 +0.039 −0.019 Alumina 0.000 0.000 0.000

indicates data missing or illegible when filed

All material cases show a decreasing trend in biological tissue reactionas a function of time elapsed post-surgery. An error occurred in theslicing of the DuPont 951 four week trial samples causing thick abnormalslices of the material. Even though this prevented accurate thicknessmeasurements, the data from the two week (14 day) implantation caseprovides a reasonable estimate of the fibrous encapsulation effects forthe DuPont 951.

After one week, the comparisons, plotted in FIG. 39, show that theHeraeus HL2000 LTCC, LCP, and silicon had the least amount ofencapsulation. LCP had the best performance out of all the materials andboth LCP and HL2000 outperformed the baseline, alumina. DuPont 943 LTCChad the highest rate of encapsulation followed by DuPont 951 that hadonly slight more fibrous tissue than the same material coated inparylene.

The tissue reaction measurements significantly decreased for allmaterials after 2 weeks, shown in FIG. 40. The percentage ofencapsulation reduction ranged from 68.1% to 96.4%. The data at 14 daysshowed that LCP and silicon caused the least amount of reaction. Somematerials (DuPont 943, HL2000, parylene coated 951, LCP, and silicon)had less encapsulation than alumina. Only DuPont 951 and ACA had moretissue reaction than the control.

The encapsulation continued to be minimal for the 28 day implantationduration, plotted in FIG. 41, decreasing in some cases from the 14 daymeasurements. LCP, parylene coated 951, and silicon had submicronencapsulation. The only material that had more reaction after the 4-weekperiod than alumina is the HL2000. The DuPont 943 and HL2000 are theonly two materials that had increased encapsulation from the 2-week to4-week periods.

Data obtained over the 28 day study allow ratings for each materialfollowing the NAMSA Good Lab Practices (GLP) protocol (NAMSA No.T1250_812). Post-surgical healing influenced encapsulation measurementsin the first week but more accurate quantifications of material-tissueinteractions are obtained at the 2 and 4-week cases. A NAMSA rating of 1represents up to 0.5 mm capsule or reaction area, which isrepresentative of all the measurements for the in-vivo studies for alltime periods. Parylene coated DuPont 951, LCP, and silicon, had nocapsule or adverse reaction after the 28 day implant duration and thushave a rating of 0. The silicon and parylene coated LTCC ratings areconsistent with that described in. This minimal tissue reaction forsilicon and parylene coated LTCC cases are also seen for the LCP.

From Table 6.2.2, there are only two points where there is astatistically helpful difference in inflammation with comparison toalumina. These occur in the Heraeus HL2000 data at 7 days, andAnisotropic Conductive Adhesive at 14 days as an over inflammationcompared to alumina.

Example 2 Implant Testing

To test in-vivo functionality, two surgical procedures have beenimplemented; one for the rabbit animal model (human IOP) and one themouse animal model. Throughout the rabbit surgical procedure, sterilesaline is applied at regular intervals to the operated eye (experimentaleye) to prevent drying. The animals are implanted in one eye (othernon-operated eye serves as histological control) with a custom,intraocular pressure monitor device in the suprachoroidal space. Afterinsertion of a lid speculum, a bridle suture can be used to isolate theeye during the procedure.

A conjunctival incision is performed posterior to the limbus, using adiagonal incision to create a flap. A 3 mm wide scleral incision to fullthickness is performed 2 mm posterior to the limbus to expose thesurface of the choroid. A blunt spatula was used to separate the scleraand choroid surface, anteriorly and posteriorly, and the pressure sensordevice is implanted into this space. If necessary, the device can besutured to the sclera to prevent migration. The sclera is closed overthe device with two 10-0 nylon sutures. If necessary, the conjunctiva isclosed with one or two 10-0 vicryl sutures. A surgical sealant isapplied to the surface of the cornea to aid in wound healing. Thesurgical sealant used will be a cyanoacrylate-based Bioglue as ourprimary anterior chamber sealant, followed by Healon if necessary. Anantibiotic, gentamicin or tobramycin was administered after woundclosure.

A stab incision of about 1 mm was made at about 3-5 mm from the corneallimbus using a 3 mm microsurgical blade, or similar device. A volume ofViscoat Viscoelastic or similar solution is used to form the anteriorchamber to allow for better manipulation of the sensor, while minimizingtrauma to the eye. (Viscoat Viscoelastic or similar solutions aresterile, non-pyrogenic, transparent viscoelastic preparations of ahighly purified, noninflammatory, high molecular weight sodiumhyaluronate or similar substances. These substances are routinely usedin anterior segment ophthalmic surgical procedures in humans. It coatsthe iris, posterior corneal surface and anterior lens capsule, whichhelps protect these tissues from injury during the surgical procedure.)The sensor is inserted through the incision, the incision sutured, andthe Viscoat Viscoelastic removed. Finally, a topical antibiotic,Vetropolycin, is applied to decrease infection risk.

As seen in the FIGS. 42 and 43, the device is properly implanted intothe eye of the rabbit and mouse respectively. Issues have arisen withthe tadpole design, as seen in FIG. 42, still using the trochanterimplantation design, whereas the implantation of the CTR device (FIG.43), was a fluid insertion without a trochanter. As the system isfurther developed, integration of a trochanter streamlines the insertionprocedure further. With this implantation procedure the mouse surgeryhas been successful in implanting the device and lighting of an LED(FIG. 45.).

Example 3 Testing of Vascularization of Eye Following Implantation ofLED Device

One eye was kept as a control and the second eye contained the implant.The results shown from the histological procedure produce results of howthe vasculature has grown due to the inflammatory responses.

The blue labeling is the vasculature and the purple is the retinatissue. It is observed that the vasculature grows toward theimplantation location, and some healing is necessary. FIG. 66 is thedepiction of the eye after one week of implantation. As the edema setsin and inflamatory response from the mouse increases, vascularization isincreasing toward the center.

A successful implantation that did not damage the cornea during surgeryis shown in FIG. 67. This will give a better understanding of theinflammatory response, since this eye was excised two weeks afterimplantation. Again using the same system, the vascularization staystoward the edge of the eye. Further analysis shows where the points ofcontact are with the implanted device as well as the implantationincision. Since the device is observed in purple, there are two aspectsto observe. These are in the top right and bottom left locations. Thetop right location is where the incision was made to implant the device,while the bottom left is the location where the tab was be interactingwith the tissue. Due to the incision procedure, the tab was caught onthe iris, causing damage. This risk needs to be mitigated in furtherdesigns of the IOP device. Further, this data can be used to make a 3Drecreation of the eye (FIG. 68). The resulting colors depict distancefrom the top of the eye (apex of the cornea). Initially the implantsubstrate sits high, while the ring and components sit lower inside theeye.

It will be understood by those having skill in the art that changes maybe made to the details of the above-described embodiments withoutdeparting from the underlying principles presented herein. For example,any suitable combination of various embodiments, or the featuresthereof, is contemplated.

Any methods disclosed herein comprise one or more steps or actions forperforming the described method. The method steps and/or actions may beinterchanged with one another. In other words, unless a specific orderof steps or actions is required for proper operation of the embodiment,the order and/or use of specific steps and/or actions may be modified.

References to approximations are made throughout this specification,such as by use of the terms “about” or “approximately.” For each suchreference, it is to be understood that, in some embodiments, the value,feature, or characteristic may be specified without approximation. Forexample, where qualifiers such as “about,” “substantially,” and“generally” are used, these terms include within their scope thequalified words in the absence of their qualifiers. For example, wherethe term “substantially planar” is recited with respect to a feature, itis understood that in further embodiments, the feature can have aprecisely planar orientation.

Reference throughout this specification to “an embodiment” or “theembodiment” means that a particular feature, structure or characteristicdescribed in connection with that embodiment is included in at least oneembodiment. Thus, the quoted phrases, or variations thereof, as recitedthroughout this specification are not necessarily all referring to thesame embodiment.

Similarly, it should be appreciated that in the above description ofembodiments, various features are sometimes grouped together in a singleembodiment, figure, or description thereof for the purpose ofstreamlining the disclosure. This method of disclosure, however, is notto be interpreted as reflecting an intention that any claim require morefeatures than those expressly recited in that claim. Rather, as thefollowing claims reflect, inventive aspects lie in a combination offewer than all features of any single foregoing disclosed embodiment.

The claims following this original written disclosure are herebyexpressly incorporated into the present written disclosure, with eachclaim standing on its own as a separate embodiment. This disclosureincludes all permutations of the independent claims with their dependentclaims. Moreover, additional embodiments capable of derivation from theindependent and dependent claims that follow are also expresslyincorporated into the present written description. These additionalembodiments are determined by replacing the dependency of a givendependent claim with the phrase “any of the preceding claims up to andincluding claim [x],” where the bracketed term “[x]” is replaced withthe number of the most recently recited independent claim. For example,for the first claim set that begins with independent claim 1, claim 3can depend from either of claims 1 and 2, with these separatedependencies yielding two distinct embodiments; claim 4 can depend fromany one of claim 1, 2, or 3, with these separate dependencies yieldingthree distinct embodiments; claim 5 can depend from any one of claim 1,2, 3, or 4, with these separate dependencies yielding four distinctembodiments; and so on. Similarly, for the second claim set that beginswith independent 10, claim 12 can depend from either of claims 10 and11, with these separate dependencies yielding two distinct embodiments;claim 13 can depend from any one of claim 10, 11, or 12, with theseseparate dependencies yielding three distinct embodiments; and so on.

Recitation in the claims of the term “first” with respect to a featureor element does not necessarily imply the existence of a second oradditional such feature or element. Elements specifically recited inmeans-plus-function format, if any, are intended to be construed inaccordance with 35 U.S.C. § 112 ¶ 6. Embodiments of the disclosure inwhich an exclusive property or privilege is claimed are defined asfollows.

What follows are further descriptions of various embodiments of thepresent invention. It is understood that these descriptions that follow(A, B, C, D) are examples of various embodiments, and are not to beconsidered as limiting statements. Following these four statements areyet further statements that describe additional aspects of any or all ofthe primary statements (A, B, C, D). It shall be understood that anyaspects or features of the below described embodiments may be combinedwith any of the embodiments or features described elsewhere herein toform still further embodiments.

A. An apparatus for measuring pressure in a media, comprising:

a sensor providing a first signal corresponding to the pressure of themedia;

an antenna for receiving and transmitting radiowaves;

a first circuit for receiving the first signal and providing a secondsignal for transmittal by said antenna;

a second circuit for receiving radiowaves from said antenna andproviding electrical power to said sensor and said second circuit.

B. An apparatus for measuring pressure in a media, comprising:

a sensor providing a first signal corresponding to the pressure of themedia;

a circuit receiving the signal and providing an output; and

an antenna for transmitting the output as a radiowave, said antennabeing fabricated from a shape memory material;

wherein said antenna does not have a free end.

C. A system for monitoring the response of an animal, comprising:

a base station including a source of radiowaves, a receiver ofradiowaves and an electronic controller including software; a sensingassembly completely implantable in the animal and adapted and configuredfor wireless communication with said base station, said sensing assemblybeing actively powered by the source and transmitting data to saidreceiver; and

a platform for supporting the animal located proximate to said receiverand proximate to said source; wherein said controller receives the datafrom said receiver, and said software provides output corresponding tothe data.

D. A method for monitoring the response of a non-human animal,comprising:

providing an actively powered sensor including an antenna fabricatedwith an elastically deformable material;

making an incision in an eye of the non-human animal;

elastically collapsing the antenna;

inserting the sensor and collapsed antenna through the incision and intoa space within the eye;

powering the sensor by radiowaves received by the antenna; and

receiving data from the sensor by radiowaves.

The apparatus of any of statements A, B, C, or D wherein the largestdimension of the apparatus is less than about ten millimeters.

The apparatus of any of statements A, B, C, or D wherein the apparatusis adapted and configured for placement in the chamber between the irisand the cornea of an eye.

The apparatus of any of statements A, B, C, or D wherein the media is afluid within an animal eye.

The apparatus of any of statements A, B, C, or D wherein the animal ishuman.

The apparatus of any of statements A, B, C, or D wherein the animal isnon-human.

The apparatus of any of statements A, B, C, or D wherein the media is afluid within a spinal cavity.

The apparatus of any of statements A, B, C, or D wherein the media is afluid within a cranial cavity.

The apparatus of any of statements A, B, C, or D wherein the media is afluid within a pliable container.

The apparatus of any of statements A, B, C, or D wherein the containeris a breast implant.

The apparatus of any of statements A, B, C, or D which further comprisesa flexible substrate, wherein said sensor, said first circuit, and saidsecond circuit are mounted to said flexible substrate.

The apparatus of any of statements A, B, C, or D wherein said sensor andsaid first circuit are in electrical communication, and thecommunication is through said substrate.

The apparatus of any of statements A, B, C, or D wherein said antenna isadapted and configured for receipt of radiowaves having a frequencygreater than about one gigahertz.

The apparatus of any of statements A, B, C, or D wherein said antennadoes not have a free end.

The apparatus of any of statements A, B, C, or D wherein said sensorincludes a capacitor the capacitance of which changes with pressure inthe media.

The apparatus of any of statements A, B, C, or D wherein said antennareceives radiowaves at a frequency and said antenna transmits the secondsignal at a harmonic of the frequency.

The apparatus of any of statements A, B, C, or D wherein said antenna isin the shape of a circular loop.

The apparatus of any of statements A, B, C, or D wherein said antennahas two ends, and each said end is attached to one of said sensor orsaid circuit.

The apparatus of any of statements A, B, C, or D wherein the materialincludes Nitinol.

The apparatus of any of statements A, B, C, or D wherein said antenna isreadily collapsible prior to implantation in a biological space of ananimal, and expands to a generally curved shape after implantation.

The apparatus of any of statements A, B, C, or D wherein said antennaexpands to a size and shape that stabilizes the apparatus at a generallyfixed location in the biological space.

The apparatus of any of statements A, B, C, or D where the geometry ofthe antenna is adapted and configured for resonance within a biologicalspace containing the media.

The apparatus of any of statements A, B, C, or D wherein the sensingassembly is a pressure sensor implanted in the eye of the animal.

The apparatus of any of statements A, B, C, or D wherein the data isprovided in real-time and the output is provided in real-time.

The apparatus of any of statements A, B, C, or D wherein said sensingassembly includes memory that is repeatedly programmable, and the memoryis programmed with information transmitted by said source.

The apparatus of any of statements A, B, C, or D wherein said source isoperates at a first frequency, and said receiver operates at a harmonicof the frequency.

The system of any of statements A, B, C, or D wherein the harmonic isthe third harmonic.

The system of any of statements A, B, C, or D wherein the platformincludes means for restraining the animal.

The apparatus of any of statements A, B, C, or D wherein the space isbetween the iris and the cornea.

The apparatus of any of statements A, B, C, or D which further comprisesexpanding the antenna to a non-collapsed shape after said inserting.

The apparatus of any of statements A, B, C, or D which further comprisestransmitting data by the antenna prior to said receiving.

The apparatus of any of statements A, B, C, or D wherein said poweringis at a frequency and said transmitting is at a different frequency.

It will be appreciated that the various apparatus and methods describedin this summary section, as well as elsewhere in this application, canbe expressed as a large number of different combinations andsubcombinations. All such useful, novel, and inventive combinations andsubcombinations are contemplated herein, it being recognized that theexplicit expression of each of these combinations is unnecessary.

APPENDIX A

When connecting circuitry that captures, or converts GHz waves, it ishelpful that this efficiency be at a high percentage. Commonly mostcircuitry is matched to a 50 ohm circuit. Therefore if two systems areconnected together and each system is seen as a 50 ohm load,theoretically the maximum efficiency is achieved when transferring powerbetween the two components. When designing the antenna and rectificationcircuitry however, these systems were not perfectly matched to 50 ohms.This mismatch (FIG. 49) can be seen between the different type ofantenna, how it is fabricated, and the loads on those systems. FIG. 49shows the imaginary and real aspects of the S11 parameter. This S11parameter describes the reflection coefficients of the components, andcan be used to determine efficiency of a circuit.

To determine efficiency, the reflection coefficient S11 or Γ is made upof a complex number (equation 8.1).

S ₁₁ =Γ=A+jB   Equation 8.1

Using the Pythagorean theorem the magnitude of the reflectioncoefficient (equation 8.2) is used to determine efficiency of a powertransfer (Equation 8.3).

$\begin{matrix}{\Gamma^{2} = {A^{2} + B^{2}}} & {{Equation}\mspace{14mu} 8.2} \\{{Eff} = {\frac{P_{act}}{P_{i\; n}} = ( {1 - {\Gamma }^{2}} )}} & {{Equation}\mspace{14mu} 8.3}\end{matrix}$

As described in Table 8.2, we see that each independent block of thesystem has different efficiencies at a given frequency. Since 2.5 GHz isused for the powering frequency, this is reflected in the results ofTable 8.3

To understand how the systems need to interact and improve efficiency,the impedance of the circuit is calculated. Since initial testing isconducted under a 50 ohm input condition, the overall impedance of eachof the circuits can be calculated using equation 8.4.

$\begin{matrix}{\Gamma = \frac{Z_{L} - Z_{0}}{Z_{L} + Z_{0}}} & {{Equation}\mspace{14mu} 8.4}\end{matrix}$

Where is the impedance of the load (antenna, rectifier, etc.) and Z₀ is50 ohms. Rearranging equation 8.4 will give an equation 8.5

$\begin{matrix}{Z_{L} = \frac{50( {1 + \Gamma} )}{( {1 - \Gamma} )}} & {{Equation}\mspace{14mu} 8.5}\end{matrix}$

Even if the efficiency as seen in Table 8.3, is low, this is withrespect to a 50 ohm load. Changes occur when the two systems arecompared with each other.

TABLE 8.3 Reflection coefficients and efficiency at 2.5 GHz 1k LoadSystem Load Real Imag Real Imag −0.244 ± 0.039 −0.533 ± 0.210 −0.122 ±0.052 −0.951 = 0.008 65.62% Eff 8.09% Eff FR4 Substrate LCP substrateReal Imag Real Imag 0.529 ± 0.001 −0.844 ± 0.000 0.592 ± 0.001 −0.718 ±0.001 0.78% Eff 13.40% Eff

Looking at the two blocks, in this case one rectifier, and one antenna,a new reflection coefficient can be determined from the two blocks.Understanding the initial conditions of

Γ_(ANT) =C+jD   Equation 8.6

Γ_(REC) =A+jB   Equation 8.7

And understanding that using equation 8.5, impedance can be determinedgiving equations 8.8 and 8.9 from equations 8.6 and 8.7.

$\begin{matrix}{Z_{ANT} = {{Y + {jZ}} = \frac{50( {C + {j\; D} + 1} )}{( {1 - ( {C + {j\; D}} )} )}}} & {{Equation}\mspace{14mu} 8.8} \\{Z_{REC} = {{W + {j\; X}} = \frac{50( {A + {j\; B} + 1} )}{( {1 - ( {A + {j\; B}} )} )}}} & {{Equation}\mspace{14mu} 8.9}\end{matrix}$

Knowing these two equations, and setting them to match to the antenna(rectifier is the load), two new equations are built from rearranging(equations 8.8 and 8.9)

Z*_(ANT)*=Y   Equation 8.10

Z* _(REC) =W+j(X−Z)   Equation 8.11

Equation 8.11 has X+Z in the imaginary due to conjugate matching of thetow systems.Then, equations 8.10 and 8.11 can be placed into the overall equationfor gamma (equation 8.4) to determine the overall efficiency usingequation 8.3 of the circuit.

$\begin{matrix}{\Gamma_{NEW} = \frac{( {W - Y + {j( {X - Z} )}} )}{( {W + Y + {j( {X - Z} )}} )}} & {{Equation}\mspace{14mu} 8.12}\end{matrix}$

Using these equations the efficiencies of both rectifiers under the fullsystem load as well as the antenna are in the single digit efficiencies.From FIG. 69 it is observed that the highest efficiencies are not in the2.5 GHz ISM band, but between 3 and 3.5 GHz. Specifically at 2.5 GHz,there are efficiencies of 0.204%, 3.74%, 0.0172, 0.328% for a comparisonof 1 kload and FR4 antenna, 1 kload and LCP antenna, full system loadand FR-4 antenna, and full system load and LCP antenna respectively.Their peaks are 71% at 3.2 GHz, 85.2% at 3.3GHz, 86.0% at 3.2 GHz, and36.0% at 3.3 GHz for the same respective comparisons.

The data obtained on the complex impedances can be further observed on asmith chart. Using that information, it is noticed that both theimpedance of the antenna and the impedance of the system lie outside the1+jω circle. Therefore, a matching network (FIG. 70) is used to increasethe efficiency between the antenna and rectifier.

FIG. 70 incorporates two extra components to match the load impedance(rectifier and system) with the source impedance (Antenna.) To determinevalues necessary for optimum transfer of power, one must first determinethe value of the components. This starts with equation 8.13.

$\begin{matrix}{\frac{1}{Z_{S}} = {{jB} + \frac{1}{R_{L} + {j( {X + {jX}_{L}} )}}}} & {{Equation}\mspace{14mu} 8.13}\end{matrix}$

Using equation 8.13, the unknown variables X and B are solved todetermine the correct inductances and capacitances. By rearrangingequation 8.13 into equation 8.14, the equation is split into terms ofreal and imaginary components (equations 8.15 and 8.16).

R _(L) J(X+X _(L))=Z _(s)−(X+X _(L))BZ _(s) +j(BR _(L) Z _(s))  Equation 8.14

(X+X _(L))BZ _(s) =Z _(s) −R _(L)   Equation 8.15

(X+X _(L))=BR_(L) Z _(s)   Equation 8.16

This gives two equations with two unknowns. Rearranging equation 8.15and placing into equation 8.16 will give one equation with one unknown.One can input data back into the second and solve for the missingparameter (equation 8.18).

$\begin{matrix}{B = {\pm \frac{\sqrt{( {Z_{S} - R_{L}} )/R_{L}}}{Z_{S}}}} & {{Equation}\mspace{14mu} 8.17} \\{X = {{\pm \sqrt{R_{L}( {Z_{S} - R_{L}} )}} - X_{L}}} & {{Equation}\mspace{14mu} 8.18}\end{matrix}$

Using the values for X and B, resulting values for capacitance andinductance are solved. As an example, using the reflection coefficientparameters from Table 8.4, the impedances of FR-4 and full system at 2.5GHz are determined using equations 8.8 and 8.9 respectively.

antenna_IMP=Z _(S)=0.416−90.3j

rectifier_IMP=Z _(L)=1.87 −44.0j

Then rewriting equations 8.8 and 8.9 as described in equation 8.10 and8.11 values are solved. Then using equations 8.17 and 8.18 the unknownvariables are solved, B=±2.12 and X=±46.3. Further, since B has minimalimpact, no component is needed for Z1, and Z2 is solved using theequation for a capacitor and obtains a value of 1.37 pF.

Using Agilent's Advanced Design Solutions (ADS), Table 8.4 is solved toshow the impedance components needed in the L-matching network. Thiscreates a conversion of power at a given powering frequency between anFR-4 antenna and the full system load.

TABLE 8.4 Components to L-match FR-4 antenna with rectifier connected tosystem load Frequency 2 2.1 2.2 2.3 2.4 2.5 2.6 2.7 2.8 2.9 3 Z1 — — — —— — — — — 0.50 — Unit — — — — — — — — — pF — Z2 0.50 0.62 0.73 0.87 1.061.36 1.90 3.10 10 — 1.03 Unit pF pF pF pF pF pF pF pF pF — nH Frequency3.1 3.2 3.3 3.4 3.5 3.6 3.7 3.8 3.9 4 Z1 — 1.09 0.61 0.75 1.11 0.85 0.560.78 0.71 0.61 Unit — pF pF pF pF pF pF pF pF pF Z2 1.68 1.84 2.08 2.081.75 1.82 2.14 1.72 1.67 1.77 Unit nH nH nH nH nH nH nH nH nH nH

1. A capacitive pressure sensor for monitoring fluid pressure within apatient, the pressure sensor comprising: a substrate; a first electrodein abutting contact with the substrate; a membrane spaced from the firstelectrode, wherein the membrane comprises: a second electrode configuredto be displaced toward the first electrode; and at least one flexiblelayer covering at least a portion of the second electrode; and adielectric region between the first electrode and the membrane, whereinthe pressure sensor has a sensitivity of no less than about 0.3 fF/mmHg.2. The pressure sensor of claim 1, wherein the second electrode isthinner than the at least one flexible layer so as to readily bend whenthe at least one flexible layer is bent toward the substrate.
 3. Thepressure sensor of claim 1, wherein the second electrode is sandwichedbetween two flexible layers.
 4. The pressure sensor of claim 1, whereineach of the substrate and the at least one flexible layer are polymeric.5. The pressure sensor of claim 1, wherein the substrate comprisesliquid crystal polymer and the at least one flexible layer comprisesparylene.
 6. The pressure sensor of claim 1, wherein the sensitivity issubstantially constant over a range of from about 0 mmHg to about 50mmHg above atmospheric pressure.
 7. The pressure sensor of claim 1,wherein a perimeter of the substrate defines an area of no greater thanabout 2 millimeters².
 8. The pressure sensor of claim 1, furthercomprising an additional flexible layer covering at least a portion ofthe first electrode, wherein the additional flexible layer and the atleast one flexible layer of the membrane comprise the same material. 9.The pressure sensor of claim 1, wherein the dielectric region comprisesa gas.
 10. The capacitive pressure sensor of claim 1,wherein themembrane has a depth of no greater than about 7 microns when themembrane is in an uncompressed position.
 11. The pressure sensor ofclaim 1, wherein the substrate is flexible so as to be able to conformto a curved surface.
 12. The pressure sensor of claim 1, wherein thepressure sensor has a sensitivity of no less than about 0.3 fF/mmHgwhether the substrate is in a flat orientation or a curved orientation.13. The pressure sensor of claim 1, wherein the first electrode isconfigured to bend in conformity with the substrate.